Switched safety protection circuit for an aimd system during exposure to high power electromagnetic fields

ABSTRACT

An energy management system that facilitates the transfer of high frequency energy induced on an implanted lead or a leadwire includes an energy dissipating surface associated with the implanted lead or the leadwire, a diversion or diverter circuit associated with the energy dissipating surface, and at least one switch disposed between the diversion circuit and the AIMD electronics for diverting energy in the implanted lead or the leadwire through the diversion circuit to the energy dissipating surface. The switch may comprise a single or multi-pole double or single throw switch. The diversion circuit may be either a high pass filter or a low pass filter.

FIELD OF INVENTION

This invention generally relates to the problem of energy induced onimplanted leads during medical diagnostic procedures such as magneticresonant imaging (MRI). Specifically, the radio frequency (RF) pulsedfield of MRI can couple to an implanted lead in such a way thatelectromagnetic forces (EMFs) are induced in the lead. The amount ofenergy that is induced is related to a number of complex factors, but ingeneral, is dependent upon the local electric field that is tangent tolead and the integral of the electric field strength along the lead. Incertain situations, these EMFs can cause currents to flow into distalelectrodes or in the electrode interface with body tissue. It has beendocumented that when this current becomes excessive, that overheating ofsaid lead or its associated electrode or overheating of the associatedinterface with body tissue can occur. There have been cases of damage tosuch body tissue which has resulted in loss of capture of cardiacpacemaking pulses, tissue damage, severe enough to result in braindamage or multiple amputations, and the like. The present inventionrelates generally to methods of redirecting said energy to otherlocations other than a distal tip electrode-to-tissue interface.

BACKGROUND OF THE INVENTION

Compatibility of probes, catheters, cardiac pacemakers, implantabledefibrillators and other types of active implantable medical deviceswith magnetic resonance imaging (MRI) and other types of hospitaldiagnostic equipment has become a major issue. If one goes to thewebsites of the major cardiac pacemaker manufacturers in the UnitedStates, which include St. Jude Medical, Medtronic and Boston Scientific(formerly Guidant), one will see that the use of MRI is generallycontra-indicated with pacemakers and implantable defibrillators. Seealso: (1) Safety Aspects of Cardiac Pacemakers in Magnetic ResonanceImaging”, a dissertation submitted to the Swiss Federal Institute ofTechnology Zurich presented by Roger Christoph Luchinger, Zurich 2002;(2) “1. Dielectric Properties of Biological Tissues: Literature Survey”,by C. Gabriel, S. Gabriel and E. Cortout; (3) “II. Dielectric Propertiesof Biological Tissues: Measurements and the Frequency Range 0 Hz to 20GHz”, by S. Gabriel, R. W. Lau and C. Gabriel; (4) “III. DielectricProperties of Biological Tissues: Parametric Models for the DielectricSpectrum of Tissues”, by S. Gabriel, R. W. Lau and C. Gabriel; and (5)“Advanced Engineering Electromagnetics, C. A. Balanis, Wiley, 1989; (6)Systems and Methods for Magnetic-Resonance-Guided InterventionalProcedures, U.S. Patent Application Publication No. US 2003/0050557,Susil and Halperin et. al, published Mar. 13, 2003; (7) MultifunctionalInterventional Devices for MRI: A Combined Electrophysiology/MRICatheter, by, Robert C. Susil, Henry R. Halperin, Christopher J. Yeung,Albert C. Lardo and Ergin Atalar, MRI in Medicine, 2002; and (8)Multifunctional Interventional Devices for Use in MRI, US 2003/0050557,and its underlying provisional application Ser. No. 60/283,725.

The contents of the foregoing are all incorporated herein by reference.

However, an extensive review of the literature indicates that MRI isindeed often used with pacemaker, neurostimulator and other activeimplantable medical device (AIMD) patients. The safety and feasibilityof MRI in patients with cardiac pacemakers is an issue of gainingsignificance. The effects of MRI on patients pacemaker systems have onlybeen analyzed retrospectively in some case reports. There are a numberof papers that indicate that MRI on new generation pacemakers can beconducted up to 0.5 Tesla (T). MRI is one of medicine's most valuablediagnostic tools. MRI is, of course, extensively used for imaging, butis also used for interventional medicine (surgery). In addition, MRI isused in real time to guide ablation catheters, neurostimulator tips,deep brain probes and the like. An absolute contra-indication forpacemaker or neurostimulator patients means that these patients areexcluded from MRI. This is particularly true of scans of the thorax andabdominal areas. Because of MRI's incredible value as a diagnostic toolfor imaging organs and other body tissues, many physicians simply takethe risk and go ahead and perform MRI on a pacemaker patient. Theliterature indicates a number of precautions that physicians should takein this case, including limiting the power of the MRI RF Pulsed field(Specific Absorption Rate—SAR level), programming the pacemaker to fixedor asynchronous pacing mode, and then careful reprogramming andevaluation of the pacemaker and patient after the procedure is complete.There have been reports of latent problems with cardiac pacemakers orother AIMDs after an MRI procedure sometimes occurring many days later.Moreover, there are a number of recent papers that indicate that the SARlevel is not entirely predictive of the heating that would be found inimplanted leads or devices. For example, for magnetic resonance imagingdevices operating at the same magnetic field strength and also at thesame SAR level, considerable variations have been found relative toheating of implanted leads. It is speculated that SAR level alone is nota good predictor of whether or not an implanted device or its associatedleadwire system will overheat.

There are three types of electromagnetic fields used in an MRI unit. Thefirst type is the main static magnetic field designated B.sub.0 which isused to align protons in body tissue. The field strength varies from 0.5to 3.0 Tesla in most of the currently available MRI units in clinicaluse. Some of the newer MRI system fields can go as high as 4 to 5 Tesla.At the recent International Society for Magnetic Resonance in Medicine(ISMRM), which was held on 5 and 6 Nov. 2005, it was reported thatcertain research systems are going up as high as 11.7 Tesla and will beready sometime in 2010. This is over 100,000 times the magnetic fieldstrength of the earth. A static magnetic field can induce powerfulmechanical forces and torque on any magnetic materials implanted withinthe patient. This would include certain components within the cardiacpacemaker itself and/or leadwire systems. It is not likely (other thansudden system shut down) that the static MRI magnetic field can inducecurrents into the pacemaker leadwire system and hence into the pacemakeritself. It is a basic principle of physics that a magnetic field musteither be time-varying as it cuts across the conductor, or the conductoritself must move within a specifically varying magnetic field forcurrents to be induced.

The second type of field produced by magnetic resonance imaging is thepulsed RF field which is generated by the body coil or head coil. Thisis used to change the energy state of the protons and elicit MRI signalsfrom tissue. The RF field is homogeneous in the central region and hastwo main components: (1) the electric field is circularly polarized inthe actual plane; and (2) the H field, sometimes generally referred toas the net magnetic field in matter, is related to the electric field byMaxwell's equations and is relatively uniform. In general, the RF fieldis switched on and off during measurements and usually has a frequencyof 21 MHz to 64 MHz to 128 MHz depending upon the static magnetic fieldstrength. The frequency of the RF pulse for hydrogen scans varies by theLamor equation with the field strength of the main static field where:RF PULSED FREQUENCY in MHz=(42.56) (STATIC FIELD STRENGTH IN TESLA).There are also phosphorous and other types of scanners wherein the Lamorequation would be different. The present invention applies to all suchscanners.

The third type of electromagnetic field is the time-varying magneticgradient fields designated B_(X), B_(Y), B_(Z), which are used forspatial localization. These change their strength along differentorientations and operating frequencies on the order of 1 kHz. Thevectors of the magnetic field gradients in the X, Y and Z directions areproduced by three sets of orthogonally positioned coils and are switchedon only during the measurements. In some cases, the gradient field hasbeen shown to elevate natural heart rhythms (heart beat). This is notcompletely understood, but it is a repeatable phenomenon. The gradientfield is not considered by many researchers to create any other adverseeffects.

It is instructive to note how voltages and electromagnetic interference(EMI) are induced into an implanted lead system. At very low frequency(VLF), voltages are induced at the input, to the cardiac pacemaker ascurrents circulate throughout the patient's body and create voltagedrops. Because of the vector displacement between the pacemaker housingand, for example, the tip electrode, voltage drop across the resistanceof body tissues may be sensed due to Ohms Law and the circulatingcurrent of the RF signal. At higher frequencies, the implanted leadsystems actually act as antennas where voltages (EMFs) are induced alongtheir length. These antennas are not very efficient due to the dampingeffects of body tissue; however, this can often be offset by extremelyhigh power fields (such as MRI pulsed fields) and/or body resonances. Atvery high frequencies (such as cellular telephone frequencies), EMIsignals are induced only into the first area of the lead system (forexample, at the header block of a cardiac pacemaker). This has to dowith the wavelength of the signals involved and where they coupleefficiently into the system.

Magnetic field coupling into an implanted lead system is based on loopareas. For example, in a cardiac pacemaker unipolar lead, there is aloop formed by the lead as it comes from the cardiac pacemaker housingto its distal tip, for example, located in the right ventricle. Thereturn path is through body fluid and tissue generally straight from thetip electrode in the right ventricle back up to the pacemaker case orhousing. This forms an enclosed area which can be measured from patientX-rays in square centimeters. Per ANSI/AAMI National Standard P069, theaverage loop area is 200 to 225 square centimeters. This is an averageand is subject to great statistical variation. For example, in a largeadult patient with an abdominal implant, the implanted loop area is muchlarger (around 400 square centimeters).

Relating now to the specific case of MRI, the magnetic gradient fieldswould be induced through enclosed loop areas. However, the pulsed RFfields, which are generated by the body coil, would be primarily inducedinto the leadwire system by antenna action. Subjected to RF frequencies,the lead itself can exhibit complex transmission line behavior.

At the frequencies of interest in MRI, RF energy can be absorbed andconverted to heat. The power deposited by RF pulses during MRI iscomplex and is dependent upon the power (Specific Absorption Rate (SAR)Level) and duration of the RF pulse, the transmitted frequency, thenumber of RF pulses applied per unit time, and the type of configurationof the RF transmitter coil used. The amount of heating also depends uponthe volume of tissue imaged, the electrical resistivity of tissue andthe configuration of the anatomical region imaged. There are also anumber of other variables that depend on the placement in the human bodyof the AIMD and its associated leadwire(s). For example, it will make adifference how much EMF is induced into a pacemaker lead system as towhether it is a left or right pectoral implant. In addition, the routingof the lead and the lead length are also very critical as to the amountof induced current and heating that would occur. Also, distal tip designis very important as the distal tip itself can heat up due to MRI RFinduced eddy currents. The cause of heating in an MRI environment istwofold: (a) RF field coupling to the lead can occur which inducessignificant local heating; and (b) currents induced between the distaltip and tissue during MRI RF pulse transmission sequences can causelocal Ohms Law heating in tissue next to the distal tip electrode of theimplanted lead. The RF field of an MRI scanner can produce enough energyto induce RF voltages in an implanted lead and resulting currentssufficient to damage some of the adjacent myocardial tissue. Tissueablation (destruction resulting in scars) has also been observed. Theeffects of this heating are not readily detectable by monitoring duringthe MRI. Indications that heating has occurred would include an increasein pacing threshold, venous ablation, Larynx or esophageal ablation,myocardial perforation and lead penetration, or even arrhythmias causedby scar tissue. Such long term heating effects of MRI have not been wellstudied yet for all types of AIMD leadwire geometries. There can also belocalized heating problems associated with various types of electrodesin addition to tip electrodes. This includes ring electrodes or padelectrodes. Ring electrodes are commonly used with a wide variety ofimplanted devices including cardiac pacemakers, and neurostimulators,and the like. Pad electrodes are very common in neurostimulatorapplications. For example, spinal cord stimulators or deep brainstimulators can include a plurality of pad electrodes to make contactwith nerve tissue. A good example of this also occurs in a cochlearimplant. In a typical cochlear implant there would be sixteen padelectrodes placed up into the cochlea. Several of these pad electrodesmake contact with auditory nerves.

Although there are a number of studies that have shown that MRI patientswith active implantable medical devices, such as cardiac pacemakers, canbe at risk for potential hazardous effects, there are a number ofreports in the literature that MRI can be safe for imaging of pacemakerpatients when a number of precautions are taken (only when an MRI isthought to be an absolute diagnostic necessity). While these anecdotalreports are of interest, they are certainly not scientificallyconvincing that all MRI can be safe. For example, just variations in thepacemaker leadwire length can significantly affect how much heat isgenerated. A paper entitled, HEATING AROUND INTRAVASCULAR GUIDEWIRES BYRESONATING RF WAVES by Konings, et al., Journal of Magnetic ResonanceImaging, Issue 12:79-85 (2000), does an excellent job of explaining howthe RF fields from MRI scanners can couple into implanted leadwires. Thepaper includes both a theoretical approach and actual temperaturemeasurements. In a worst-case, they measured temperature rises of up to74 degrees C. after 30 seconds of scanning exposure. The contents ofthis paper are incorporated herein by reference.

The effect of an MRI system on the function of pacemakers, ICDs,neurostimulators and the like, depends on various factors, including thestrength of the static magnetic field, the pulse sequence, the strengthof RF field, the anatomic region being imaged, and many other factors.Further complicating this is the fact that each patient's condition andphysiology is different and each manufacturer's pacemaker and ICDdesigns also are designed and behave differently. Most experts stillconclude that MRI for the pacemaker patient should not be consideredsafe.

It is well known that many of the undesirable effects in an implantedlead system from MRI and other medical diagnostic procedures are relatedto undesirable induced EMFs in the lead system and/or RF currents in itsdistal tip (or ring) electrodes. This can lead to overheating of bodytissue at or adjacent to the distal tip.

Distal tip electrodes can be unipolar, bipolar and the like. It is veryimportant that excessive current not flow at the interface between thelead distal tip electrode and body tissue. In a typical cardiacpacemaker, for example, the distal tip electrode can be passive or of ascrew-in helix type as will be more fully described. In any event, it isvery important that excessive RF current not flow at this junctionbetween the distal tip electrode and for example, myocardial or nervetissue. This is because tissue damage in this area can raise the capturethreshold or completely cause loss of capture. For pacemaker dependentpatients, this would mean that the pacemaker would no longer be able topace the heart. This would, of course be life threatening for apacemaker dependent patient. For neurostimulator patients, such as deepbrain stimulator patients, the ability to have an MRI is equallyimportant.

A very important and life-threatening problem is to be able to controloverheating of implanted leads during an MRI procedure. A novel and veryeffective approach to this is to install parallel resonant inductor andcapacitor bandstop filters at or near the distal electrode of implantedleads. For cardiac pacemaker, these are typically known as the tip andring electrodes. One is referred to U.S. Pat. No. 7,363,090; US2007/0112398 A1; US 2008/0071313 A1; US 2008/0049376 A1; US 2008/0024912A1; US 2008/0132987 A1; and US 2008/0116997 A1, the contents of all ofwhich are incorporated herein. Referring now to US 2007/0112398 A1, theinvention therein relates generally to L-C bandstop filter assemblies,particularly of the type used in active implantable medical devices(AIMDs) such as cardiac pacemakers, cardioverter defibrillators,neurostimulators and the like, which raise the impedance of internalelectronic or related wiring components of the medical device atselected frequencies in order to reduce or eliminate currents inducedfrom undesirable electromagnetic interference (EMI) signals.

U.S. Pat. No. 7,363,090 and US 2007/0112398 A1 show resonant L-Cbandstop filters to be placed at the distal tip and/or at variouslocations along the medical device leadwires or circuits. These L-Cbandstop filters inhibit or prevent current from circulating at selectedfrequencies of the medical therapeutic device. For example, for an MRIsystem operating at 1.5 Tesla, the pulse RF frequency is 64 MHz, asdescribed by the Lamour Equation for hydrogen. The L-C bandstop filtercan be designed to resonate at or near 64 MHz and thus create a highimpedance (ideally an open circuit) in the lead system at that selectedfrequency. For example, the L-C bandstop filter, when placed at thedistal tip electrode of a pacemaker lead, will significantly reduce RFcurrents from flowing through the distal tip electrode and into bodytissue. The L-C bandstop filter also reduces EMI from flowing in theleadwires of a pacemaker, for example, thereby providing added EMIprotection to sensitive electronic circuits.

Electrically engineering a capacitor in parallel with an inductor isknown as a bandstop filter or tank circuit. It is also well known thatwhen a near-ideal L-C bandstop filter is at its resonant frequency, itwill present a very high impedance. Since MRI equipment produces verylarge RF pulsed fields operating at discrete frequencies, this is anideal situation for a specific resonant bandstop filter. Bandstopfilters are more efficient for eliminating one single frequency thanbroadband filters. Because the L-C bandstop filter is targeted at thisone frequency, it can be much smaller and volumetrically efficient.

A major challenge for designing an L-C bandstop filter for human implantis that it must be very small in size, biocompatible, and highlyreliable. Coaxial geometry is preferred. The reason that coaxial ispreferred is that implanted leads are placed at locations in the humanbody primarily by one of two main methods. These include guide wire leadinsertion. For example, in a cardiac pacemaker application, a pectoralpocket is created. Then, the physician makes a small incision betweenthe ribs and accesses the subclavian vein. The pacemaker leadwires arestylus guided/routed down through this venous system through thesuperior vena cava, through the right atrium, through the tricuspidvalve and into, for example, the right ventricle. Another primary methodof implanting leads (particularly for neurostimulators) in the humanbody is by tunneling. In tunneling, a surgeon uses special tools totunnel under the skin and through the muscle, for example, up throughthe neck to access the Vagus nerve or the deep brain. In bothtechniques, it is very important that the leads and their associatedelectrodes at the distal tips be very small. US 2007/0112398 A1 solvesthese issues by using miniature coaxial or rectilinear capacitors thathave been adapted with an inductance element to provide a parallel L-Cbandstop filter circuit.

Prior art capacitors used in design of bandstop filters typicallyconsist of ceramic discoidal feedthrough capacitors and also singlelayer and multilayer tubular capacitors and multilayer rectangularcapacitors, and thick-film deposited capacitors. US 2007/0112398 A1shows design methodologies to adapt all of these previous tubular,feedthrough or rectangular technologies to incorporate a parallelinductor in novel ways. It will be obvious to those skilled in the artthat a number of other capacitor technologies can be adapted. Thisincludes film capacitors, glass capacitors, tantalum capacitors,electrolytic capacitors, stacked film capacitors and the like.

As previously mentioned, the value of the capacitance and the associatedparallel inductor can be adjusted to achieve a specific resonantfrequency (SRF). The bandstop filters described in US 2007/0112398 A1can be adapted to a number of locations within the overall implantablemedical device system. That is, the L-C bandstop filter can beincorporated at or near any part of the medical device implanted leadsystem or at or adjacent to the distal tip electrodes. In addition, theL-C bandstop filter can be placed anywhere along the implanted leadsystem.

The L-C bandstop filters are also designed to work in concert with anEMI filter which is typically used at the point of leadwire ingress andegress of the active implantable medical device. For example, see U.S.Pat. No. 5,333,095; U.S. Pat. No. 5,905,627; U.S. Pat. No. 5,896,627;and U.S. Pat. No. 6,765,779, the contents of all being incorporatedherein by reference. All four of these documents describe low pass EMIfilter circuits. Accordingly, the L-C bandstop filters, as described inU.S. Pat. No. 7,393,090, are designed to be used in concert with suchlow pass filters.

When the value of a hermetic feedthrough filter capacitor is too high,the leading edge of MRI gradient pulse sequences can create an R-Ccharging circuit. As the feedthrough capacitor charges up this voltagefall can create one of two problems. First, the voltage induced on theleadwire system could directly capture the heart thereby creating adangerously rapid heart rate which could then result in a dangerousventricular arrhythmia. For example, ventricular fibrillation can resultin sudden death. Another problem associated with too high of a value ofa feedthrough capacitor at the input to the AIMD is that this R-Ccharging circuit can cause pulses to appear at the input sense amplifier(such as a cardiac pacemaker) such that the pacemaker would oversense orfalsely interpret this input as a normal heartbeat. In certain casesthis can cause a demand pacemaker to inhibit (stop pacing). For apacemaker dependent patient this can lead to systole and be immediatelylife threatening. Accordingly, it is desirable for magnetic resonancecompatibility to keep the value of the feedthrough capacitor relativelylow (in the order of 1000 picofarads). On the other hand, in order toadequately protect AIMD device electronics from the powerful RF pulsefield of MRI, we have a trade off in that it would be desirable to havethe hermetic feedthrough capacitor be as large as value as possible (inthe order of 4,000 to 6,000 picofarads).

When one performs MRI testing on an active implantable medical device(AIMD) with its associated lead system, one first establishes acontrolled measurement. That is, with worst-case MRI equipment settingsand a worst-case location within the MRI bore, and a worst-case leadconfiguration, one can measure heating using fiber optic probes at thedistal electrodes. Temperature rises of 30 to over 60 degrees C. havebeen documented. When one takes the same control lead and placesminiature bandstop filters in accordance with U.S. Pat. No. 7,363,090 orUS 2007/0112398 A1, one finds that the distal electrodes aresubstantially cooled. In fact, in many measurements made by theinventors, temperature rises of over 30 degrees C. have been reduced toless than 3 degrees C. However, a secondary problem has been discovered.That is, the implanted lead acts very much as like a transmission line.When one creates a very high impedance at the distal electrode to tissueinterface by installation of a resonant bandstop filter as described inU.S. Pat. No. 7,038,900 and as further described in US 2007/0112398 A1,there is created an almost open circuit which is the equivalent of anunterminated transmission line. This causes a reflection of MRI inducedRF energy back towards the AIMD (for example, toward the pacemakerhousing). This energy can be reflected back and forth resulting intemperature rises along the lead. In some cases, the inventors havemeasured temperature rises immediately proximal to the bandstop filters,which is undesirable.

Accordingly, there is a need for controlling the induced energy inimplanted lead system. This may be accomplished by taking a systemapproach and carefully balance the filtering needs. Moreover, there is aneed for novel tuned RE diverting circuits coupled to one or more energyor heat dissipation surfaces, which are frequency selective and areconstructed of passive components. Such circuits are needed to preventMRI induced energy from reaching the distal tip electrode or itsinterface with body tissue. By redirecting said energy to an energydissipation surface distant from the distal electrodes, this minimizesor eliminates hazards associated with overheating of said lead and/orits distal electrodes during diagnostic procedures, such as MRI.Frequency selective diverter circuits are needed which decouple andtransfer energy which is induced onto implanted leads from the MRIpulsed RF field to an energy dissipating surface. In this regard, anovel system is needed which can utilize the conductive housing (can) ofthe AIMD itself as the energy dissipation surface. A switched divertercircuit would be beneficial in such a system for minimizing heating ofan implanted lead in a high power electromagnetic field environment. Thepresent invention fulfills these needs and provides other relatedadvantages.

SUMMARY OF THE INVENTION

The present invention includes an overall energy management systemcapable of controlling the energy induced in implanted leads from the RFpulsed field of MRI scanners. The term “implanted lead” as used hereinincludes permanently implanted leads or long-term implanted leads suchas those that might be associated with an active implantable medicaldevice such as a cardiac pacemaker. However, “implanted lead” or “lead”as used herein also includes temporary implants such as those fromprobes or catheters. For example, it is becoming increasingly importantto perform real time mapping and catheter ablation while using MRI forreal-time imaging of scar tissue. The term “implanted lead” as usedherein can also include temporary implanted leads from loop recorders,probes or catheters or even implanted leads that are attached to anexternal device such as externally worn spinal cord or pain controlstimulator.

More particularly, the present invention relates to a switched safetyprotection circuit for an AIMD system during exposure to high powerelectromagnetic fields. The switched safety protection circuit comprises(1) an active implantable medical device (AIMD) including AIMDelectronics, (2) an implanted lead or leadwire associated with the AIMDelectronics, (3) an energy dissipating surface associated with theimplanted lead or the leadwire, (4) a diversion circuit associated withthe energy dissipating surface, and (5) at least one switch disposedbetween the diversion circuit and the AIMD electronics, for divertingenergy.

In some embodiments, the diversion circuit is permanently conductivelyconnected between the implanted lead or the leadwire and the energydissipating surface. In others, a switch is associated with thediversion circuit and disposed between the lead or the leadwire and theenergy dissipating surface. In this case, the switch associated with thediversion circuit may be disposed between the implanted lead or theleadwire and the diversion circuit. Alternatively, the switch associatedwith the diversion circuit may be disposed between the diversion circuitand the energy dissipating surface. The switch associated with thediversion circuit may comprise a single or multi-pole single ormulti-throw switch. Said switch may comprise a MEMS switch, a mechanicalswitch, a reed switch, an electronic switch, a programmable switch, anautomatically actuated switch, a Hall Effect switch, or a field effecttransistor (FET) switch.

At least one switch may electrically open the implanted lead or theleadwire when diverting energy in the implanted lead or the leadwirethrough the diversion circuit to the energy dissipating surface, andsimultaneously conductively couple the AIMD electronics to the energydissipating surface. Here again, the switch may comprise a single or amulti-pole double-throw switch.

The AIMD electronics may include non-linear circuit elements,over-voltage circuit elements, transient voltage suppression diodes,Transorbs, AIMD sensing circuits and therapy delivery circuits.

The implanted lead or the leadwire has impedance characteristics at aselected RF frequency or frequency band. The diversion circuit likewisehas impedance characteristics least partially tuned to the implantedlead's or the leadwire's impedance characteristics. The selected RFfrequency or frequency band may comprise an MRI frequency or a range ofMRI frequencies. The diversion circuit also typically has a reactanceand is vectorally opposite to the characteristic reactance of theimplanted lead. Moreover, the diversion circuit typically has acapacitive reactance generally equal and opposite to the characteristicinductive reactance of the implanted lead. The capacitive reactance andthe inductive reactance each have a resistor component.

The diversion circuit may comprise a low pass filter comprising acapacitor, an inductor, a Pi filter, a T an LL filter, or an “n” elementfilter. Moreover, the diversion circuit may comprise a unipolar ormultipolar feedthrough capacitor, at least one series resonant L-C trapfilter, or a plurality of L-C trap filters resonant respectively atdifferent MRI frequencies.

The energy dissipating surface may comprise a housing for an activeimplantable medical device (AIMD). The AIMD may comprise an implantablehearing device, a neurostimulator, a brain stimulator, a cardiacpacemaker, a left ventricular assist device, an artificial heart, a drugpump, a bone growth stimulator, a urinary incontinence device, a spinalcord stimulator, an anti-tremor stimulator, an implantable cardioverterdefibrillator, a congestive heart failure device, or a cardioresynchronization therapy device.

The diversion circuit may be disposed within the housing for the AIMD,and further may comprise a short to the housing. Alternatively, thediversion circuit may be disposed within a header block for the AIMD.

An EMI shield may be conductively coupled to the housing and coaxiallyextend about the lead or the leadwire in non-conductive relation. Thediversion circuit may be conductively coupled to the EMI shield.

The diversion circuit may also comprise a high pass filter whichprevents low frequency gradient field-induced energy in the implantedlead or the leadwire from passing through the diversion circuit to theenergy dissipating surface. The high pass filter may comprise a resistorin series with a capacitor or an L-C trap filter.

The implanted lead typically has a length extending between and to aproximal end and a tissue-stimulating or biological-sensing electrode ator near a distal tip end. The diversion circuit may be disposed at oradjacent to the proximal end of the implanted lead or in a proximal leadconnector.

An impedance circuit may be associated with the diversion circuit, forraising the high-frequency impedance of the implanted lead. Theimpedance circuit may comprise an inductor, a bandstop filter and/or aresistor.

The implanted lead may comprise at least a portion of a probe or acatheter, wherein the energy dissipating surface may comprise at a leasta portion of a handle for the probe or the catheter.

The implanted lead may comprise at least a pair of leads each having alength extending between and to a proximal end and a tissue-stimulatingor biological-sensing electrode at a distal tip end. The diversioncircuit may couple each of the leads to the energy dissipating surface,or the diversion circuit may be coupled between the pair of leads.

The energy dissipating surface may comprise a material capable of beingvisualized during a magnetic resonance scan, and include a biomimeticcoating.

Further, the diversion circuit may include at least one non-linearcircuit element such as a transient voltage suppressor or a diode.

Other features and advantages of the present invention will becomeapparent from the following more detailed description, taken inconjunction with the accompanying drawings which illustrate, by way ofexample, the principles of the invention.

BRIEF DESCRIPTION OF THE DRAWINGS

The accompanying drawings illustrate the invention. In such drawings:

FIG. 1 is a wire-formed diagram of a generic human body showing a numberof exemplary implanted medical devices;

FIG. 2 is a diagrammatic view of a typical probe or catheter;

FIG. 3 is an electrical diagrammatic view of the interior of the proberor catheter of FIG. 2;

FIG. 4 is an electrical diagrammatic view of the structure shown in FIG.3, with a general impedance element connected between leadwires;

FIG. 5 is an electrical diagrammatic view similar to FIG. 4,illustrating a capacitor representing a frequency dependent reactiveelement between the leadwires;

FIG. 6 is a view similar to FIG. 5, wherein the general reactanceelement has been replaced by a capacitor in series with an inductor;

FIG. 7 is a view similar to FIGS. 4-6, showing the addition of seriesfrequency selective reactances;

FIG. 8 is similar to FIG. 3, showing a low frequency model of thecatheter and associated leads described in FIG. 2;

FIG. 9 is a view similar to FIGS. 3-8, illustrating how the distal ringsare electrically isolated at a high frequency;

FIG. 10 is a view similar to FIGS. 3-9, showing the addition of seriesinductor components added to the frequency selective elements 20;

FIG. 11 is similar to FIGS. 3-10, illustrating frequency selectiveelements which incorporate parallel resonant inductor and capacitorbandstop filters;

FIG. 12 is a perspective and somewhat schematic view of a prior artactive implantable medical device (AIMD) including a leadwire directedto the heart of a patient;

FIG. 13 is a schematic illustration of a bipolar leadwire system with adistal tip and ring typically as used with a cardiac pacemaker;

FIG. 14 is a schematic illustration of a prior art single chamberbipolar cardiac pacemaker lead showing the distal tip and the distalring electrodes;

FIG. 15 is an enlarged, fragmented schematic view taken generally alongthe line 15-15 of FIG. 14, illustrating placement of bandstop filtersadjacent to the distal tip and ring electrodes;

FIG. 15 is a tracing of an exemplary patient X-ray showing an implantedpacemaker and cardioverter defibrillator and corresponding leadwiresystem;

FIG. 17 is a line drawing of an exemplary patient cardiac X-ray of abi-ventricular leadwire system;

FIG. 18 is a chart showing the calculation of the frequency of resonancefor a parallel L-C tank circuit of FIG. 15;

FIG. 19 is a graph showing impedance versus frequency for the idealparallel tank circuit of FIG. 15;

FIG. 20 illustrates the equation for the impedance for the inductor inparallel with the capacitor;

FIG. 21 gives the equations for inductive reactance X_(L) and capacitivereactance X;

FIG. 22 is an enlarged schematic illustration of the area indicated byLine 22-22 in FIG. 15, showing details of the bipolar pacemaker leadwiresystem;

FIG. 23 is similar to FIG. 22, but depicts an active fixation tip for abipolar pacemaker leadwire system;

FIG. 24 is similar to FIG. 22, except that the twisted or coaxialelectrode wires have been straightened out;

FIG. 25 is similar to FIG. 24 and incorporates electrical featuresdiscussed in FIGS. 2-11;

FIG. 26 is similar to a portion of FIGS. 24 and 25, and depicts an L-Ctrap filter coupled between a distal tip electrode wire and acylindrical ring electrode;

FIG. 27 is a line drawing of a human heart with cardiac pacemaker dualchamber bipolar leads shown in the right ventricle and the right atrium;

FIG. 28 is a schematic diagram illustration of an energy dissipatingsurface in spaced relation with tip and ring electrodes;

FIG. 29 is a schematic diagram depicting a typical quad polarneurostimulation lead system;

FIG. 30 is a somewhat schematic side view of the human head with a deepbrain stimulation electrode shaft assembly implanted therein;

FIG. 31 is an enlarged sectional view corresponding generally with theencircled region 31-31 of FIG. 30;

FIG. 32 is a further enlarged and somewhat schematic view correspondinggenerally with the encircled region 32-32 of FIG. 31;

FIG. 33 is an enlarged and somewhat schematic view correspondinggenerally with the encircled region 33-33 of FIG. 31;

FIG. 34 is a sectional view of an hermetically sealed electrode assemblydesigned for contact with body fluid;

FIG. 35 is a perspective sectional view of a housing portion of thesealed electrode assembly of FIG. 34;

FIG. 36 is an enlarged sectional view corresponding generally with theencircled region 36-36 of FIG. 35, and illustrating the principle ofincreasing the surface area of the energy dissipating surface;

FIG. 37 is a schematic circuit diagram corresponding with the sealedelectrode assembly of FIG. 34;

FIG. 38 is a perspective view of an exemplary monolithic capacitor foruse in the circuit of FIG.

FIG. 39 is a perspective view of an exemplary unipolar feedthroughcapacitor for use in the circuit of FIG. 37;

FIG. 40 is a sectional view similar to FIG. 34 and depicts analternative embodiment wherein an inductor element is wound or printedabout a central mandrel;

FIG. 41 is a sectional view similar to FIGS. 34 and 40, but illustratesa further alternative embodiment of the invention with alternative meansfor decoupling signals from a leadwire to an energy dissipating surface;

FIG. 42 is a schematic circuit diagram corresponding with the sealedelectrode assembly of FIG. 41;

FIG. 43 is an attenuation versus frequency chart for various types oflow pass filters;

FIG. 44 shows schematic circuit diagrams for different types of low passfilters charted in FIG. 43;

FIG. 45 is a schematic circuit diagram illustrating an L-C trap filter;

FIG. 46 depicts a resonant frequency equation for the L-C trap filter ofFIG. 45;

FIG. 47 is an impedance versus frequency chart for the L-C trap filterof FIG. 35;

FIG. 48 is a sectional view similar to FIGS. 34, 40 and 41, but showsstill another alternative embodiment of the invention for decoupling RFsignals from an electrode leadwire;

FIG. 49 is a schematic circuit diagram corresponding with the sealedelectrode assembly of FIG. 48;

FIG. 50 illustrates a typical chip inductor for use in the sealedelectrode assembly of FIG. 48;

FIG. 51 illustrates a typical chip capacitor for use in the sealedelectrode assembly of FIG. 48;

FIG. 52 is an impedance versus frequency chart for the dual L-C trapfilter embodiment of FIG. 48;

FIG. 53 is a schematic representation of an implantable medical devicebipolar leadwire system;

FIG. 54 is an enlarged and somewhat schematic sectional view takengenerally on the line 54-54 of FIG. 53;

FIG. 55 is an isometric view of a bipolar feedthrough capacitor for usein the device of FIGS. 53-54;

FIG. 56 is a schematic circuit diagram corresponding with the embodimentshown in FIGS. 53-54;

FIG. 57 is a schematic circuit diagram illustrating a bipolar leadassembly with distal tip and ring electrodes shown at a suitabledistance from an energy dissipation surface;

FIG. 58 is a schematic circuit diagram similar to FIG. 57, except that apair of capacitors are used;

FIG. 59 is a schematic circuit diagram illustrating a bandstop filtermodified to include a pair of diodes in a parallel or back-to-backconfiguration;

FIG. 60 is a schematic circuit diagram similar to FIG. 58, except thattransient voltage suppressors are installed in parallel relation witheach of the bandstop filter elements;

FIG. 61 is a schematic circuit diagram depicting a general filterelement constructed in accordance with any one of the embodiments shownand described herein, wherein the filter element is coupled between thedistal and proximal ends of a leadwire or the like, for dissipating RFenergy or heat to an adjacent energy dissipating surface;

FIG. 62 is a schematic circuit diagram similar to FIG. 61, but showingalternative design considerations;

FIG. 63 depicts in somewhat schematic form a probe or catheterconstructed in accordance with the present invention;

FIG. 64 is an illustration similar to FIG. 63, illustrating analternative embodiment wherein the energy dissipating surface has beenconvoluted so that its surface area has been increased;

FIG. 65 is similar to FIG. 64, except that instead of convolutions, finshave been added to the energy dissipating surface;

FIG. 66 is similar to FIGS. 64 and 65, except that the energydissipating surface has its surface area increased through varioussurface roughening processes;

FIG. 67 is an enlarged, fragmented sectional view taken along the line67-67 from FIG. 66, illustrating a roughened surface formed through, forexample, plasma or chemical etching, or the like;

FIG. 68 is a view similar to FIG. 67, and illustrates the use of carbonnanotubes or fractal coatings to increase the surface area of the energydissipating surface;

FIG. 69 is an illustration of a steerable catheter;

FIG. 70 is an enlarged section view taken generally along the line 70-70from FIG. 69;

FIG. 71 is a schematic view of a probe or catheter similar to FIG. 63,except that the number of individual energy dissipating surfaces havebeen provided in distinct and spaced-apart segments;

FIG. 72 is a fragmented top plan view of an exemplary paddle electrodeembodying the present invention;

FIG. 73 is a bottom plan view of the paddle electrode shown in FIG. 72;

FIG. 74 is an enlarged sectional view taken generally along the line74-74 in FIG. 72;

FIG. 75 is a top plan view of a different type of paddle lead structurein comparison with that shown in FIGS. 72-74;

FIG. 76 is an enlarged electrical schematic view taken generally of thearea indicated by the line 76-76 in FIG. 75;

FIG. 77 is a schematic illustration similar to FIG. 30, showing use of atethered energy dissipating surface in accordance with the presentinvention;

FIG. 78 is an enlarged sectional view of the area indicated by the line78-78 in FIG. 77;

FIG. 79 is an enlarged, somewhat schematic illustration of thecomponents found within the area designated by the line 79-79 in FIG.78;

FIG. 80 is an overall outline drawing showing a cardiac pacemaker withendocardial leadwires implanted into a human heart;

FIG. 81 is an illustration of an AIMD similar to FIG. 12, illustratingthe use of variable impedance elements in connection with a leadwirewithin the housing of the AIMD;

FIG. 82 is a schematic illustration of the structure shown in FIG. 81,showing use of variable impedance elements on leads that ingress andegress the AIMD;

FIG. 83 is a schematic illustration showing that a variable impedanceelement can be a capacitor element;

FIG. 84 is a schematic illustration similar to FIG. 83, showing that thevariable impedance element can be a feedthrough capacitor element;

FIG. 85 is a schematic illustration similar to FIG. 84, showing use of acapacitor element in parallel with the L-C trap filter;

FIG. 86 is similar to FIG. 81 with emphasis on the series variableimpedance element 118;

FIG. 87 illustrates that the variable impedance element 118 can be aninductor;

FIG. 88 illustrates that the variable impedance element 118 can be anL-C bandstop filter;

FIG. 89 is a schematic illustration of a unipolar lead system for anAIMD;

FIG. 90 is an illustration similar to FIG. 89, wherein an L-C trapfilter has been placed inside an abandoned lead cap assembly;

FIG. 91 is another illustration similar to FIG. 89, wherein thefrequency selective components comprise capacitive elements;

FIG. 92 is another illustration similar to FIG. 89, illustrating thesimplest arrangement where the leadwire is shorted to the housing of theAIMD;

FIG. 93 is another illustration similar to FIGS. 89 and 91, wherein thecapacitance value C has been selected such that the capacitive reactancewill be equal and opposite to the inductive reactance of the implantedlead;

FIG. 94 is a schematic illustration similar to FIG. 89, illustrating anovel switch;

FIG. 95 is similar to FIG. 94, wherein the novel switch is showndisposed between the leadwires;

FIG. 96 is similar to FIG. 94, wherein two switches are employed;

FIG. 97 is similar to FIG. 96, wherein the position of the two switchesare changed;

FIG. 98 is similar to FIGS. 96-97 except that switches have beenreplaced by a single switch which can be either a single or multipoledouble pole switch;

FIG. 99 illustrates the switch in FIG. 98 can be disposed within anenergy dissipating surface located anywhere along an implanted lead;

FIG. 100 is similar to FIG. 98, wherein a capacitor is connected betweenthe switch point and the energy dissipating surface;

FIG. 101 is an illustration of an X-ray tracing of an implanted cardiacpacemaker in a patient;

FIG. 102 is similar to FIG. 100, where the circuit is a high frequencymodel;

FIG. 103 is similar to FIG. 100, where the circuit is a low frequencymodel;

FIG. 104 shows the application of the switch of FIG. 98 to a circuitboard or housing of an AIMD;

FIG. 105 similar to FIG. 104, except that the short has been replaced bya frequency variable impedance element;

FIG. 106 is similar to FIGS. 104-105, except that the diverter elementis shown as a capacitor;

FIG. 107 is similar to FIG. 106, except that a resistor has been addedin series with the diverter capacitor;

FIG. 108 is similar to FIGS. 104-107, except that the diverter elementis an RLC trap filter;

FIG. 109 is an illustration of connecting a flex cable to leadwires thatpass through the hermetic terminal in non-conductive relationship;

FIG. 110 is a cross-sectional view taken generally from the attachmentof the flex cable to the hermetic terminal previously described in FIG.109;

FIG. 111 is taken from section 111-111 of FIG. 110 and illustrates theinternal circuit traces;

FIG. 112 is taken from section 112-112 of FIG. 110 and illustrates oneof a pair of coaxially surrounding shields disposed about the circuittraces;

FIG. 113 is taken from section 113-113 of FIG. 110 and illustrates analternative to the layer previously described in FIG. 111;

FIG. 114 illustrates the flex cable arrangement of FIGS. 109-113 nowconnected to a circuit board or substrate;

FIG. 115 is similar to FIG. 81, except that it shows optional locationsfor the frequency selective diverter elements and the frequencyselective impeder elements;

FIG. 116 is similar to FIGS. 98 and 100, illustrating an alternativeembodiment; and

FIG. 117 is similar to FIG. 116 except that the switch is a single-poledouble throw switch.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

As shown in the drawings for purposes of illustration, the presentinvention resides in a tuned energy balanced system including a switcheddiverter circuit, for minimizing heating of an implanted lead in a highpower electromagnetic field environment. The tuned energy balancedsystem comprises an implanted lead having impedance characteristics at aselected RF frequency or frequency band, and an energy dissipatingsurface associated with the implanted lead. An energy diversion circuitconductively couples the implanted lead to the energy dissipatingsurface. The diversion circuit comprises one or more passive electronicnetwork components whose impedance characteristics are at leastpartially tuned to the implanted lead's impedance characteristics, tofacilitate transfer to the energy dissipating surface of high frequencyenergy induced on the implanted lead at the selected RF frequency orfrequency band.

More particularly, the present invention resides in a switched divertercircuit comprising at least one switch for diverting energy in theimplanted lead or a leadwire through the diversion circuit to the energydissipating surface. In alternate embodiments, the switch is disposedbetween the implanted lead or the leadwire and the diversion circuit, orit is situated such that it electrically opens the implanted lead or theleadwire when diverting energy in the implanted lead or the leadwirethrough the diversion circuit to the energy dissipating surface. Theswitch may comprise a single or a multi-pole double throw or singlethrow switch.

In one preferred embodiment, the invention resides in a combination ofbandstop filters placed at or near the distal electrode-to-tissueinterface of implanted leads such bandstop filters are best used incombination with the novel frequency selective diverter circuits of thepresent invention decouple energy induced on the implanted leads at afrequency or frequency band of an interest to an energy dissipatingsurface such as the AIMD housing. This can happen to a certain extent,when the AIMD already has a low-pass filter capacitor at its point ofleadwire ingress or egress. As previously mentioned, feedthroughfiltered capacitors are well known in the art. For example, reference ismade to U.S. Pat. Nos. 4,424,551; 5,333,095; 6,765,779 and the like.Feedthrough capacitors, of course, are well known in the prior art andare used as EMI filters. For example, in a cardiac pacemakerapplication, it is very common that feedthrough-type filter capacitors,at the point of leadwire ingress/egress to the pacemaker housing, wouldbe used to decouple various signals from cell phones and other emitterstypically found in the patient environment. As will be described herein,it is very important that the value of a low pass filter capacitor becarefully selected such that the capacitive reactance is approximatelyequal to the characteristic inductive reactance of the implanted leadsystem. Certain implanted leads may have a characteristic impedance thatincludes capacitive reactance. In this case the novel diverter circuitsof the present invention would include inductive elements in order tocancel the capacitive reactance of the implanted lead. However, thereare many types of AIMDs that do not use feedthrough filters. These AIMDsgenerally do not have sense circuits and are therefore inherently muchless susceptible to electromagnetic interference. Examples include manytypes of neurostimulators, including pain control stimulators, bladdercontrol stimulators, deep brain stimulators, cochlear implants and thelike.

Accordingly, when bandstop filters are installed at or near the distalelectrode of an implanted lead, the RF energy induced by the MRI pulsefield is prevented from flowing into body tissues and thereby beingdissipated. However, when bandstop filters are used, that energy stillresides in the lead system. In other words, by preventing this inducedenergy from flowing to sensitive tissues at distal electrode interfaces,a great deal has been accomplished; however, it is still important tocarefully dissipate the remaining energy that's trapped in the leadsystem. The most efficient way to do this is to use the metallic housingof the AIMD. One type of frequency selective network, is of course, theprior art feedthrough capacitors that were used for EMI filters.However, to provide optimal decoupling, one has to refer to the maximumpower transfer theorem. When one has an ideal source, consisting of avoltage source and a series impedance, this is known as a TheveninEquivalent Circuit. It is well known in electrical engineering that totransfer maximum power to a load that the load impedance must be equalto the source impedance. If the source impedance is completelyresistive, for example, 50 ohms, then to transfer maximum power, theload impedance would have to be 50 ohms. When the source impedance isreactive, then to transfer maximum power to another location, the loadimpedance should have the opposite sign of reactance and the sameimpedance and resistance. Referring to a typical implanted lead system,the implanted leads typically appear inductive. Accordingly, having acapacitive load at the point of leadwire ingress/egress to the AIMDhousing, one has at least some cancellation of these imaginary impedancefactors. In electrical engineering, the inductance of the lead would bedenoted by +jωL. The impedance of the capacitor, on the other hand, is a−j/ωC term. In the present invention, it's important to know theinductance property of the implanted lead system, so that an optimalvalue of capacitance between the AIMD housing and ground can be selectedsuch that the +J component is nearly or completely canceled by theappropriate −J component of the capacitor.

However, for devices that have sense circuits (such as cardiacpacemakers and implantable cardioverter defibrillators), the typicallead impedance values result in a capacitance at the point of leadwireingress that is too low for effective AIMD EMI protection. That is, theamount of inductance in the implanted lead system is relatively lowwhich results in a cancellation capacitor which is also relatively low.This result in a capacitance value at the point of leadwire ingress andegress that is too low to effectively attenuate a broad range of EMIfrequencies, including cellular telephones all the way down to, forexample, 13.56 MHz RFID systems. In other words, one has, for someAIMDs, a trade-off that is simply unacceptable. On the one hand, formaximal MRI energy dissipation from the lead system, one would want thecapacitance value to be equal and opposite in value to the inductance ofthe lead system, which would result in a capacitance value that would betoo low (in the order of a few hundred picofarads). However, for optimalEMI filtering, one desires a filtered capacitor in the area of a fewthousand picofarads. In other words, there is an order of magnitudeproblem here. As previously mentioned, this problem does not exist forAIMDs that do not employ feedthrough capacitors (in general, AIMDs thatdo not have sense circuits).

One does not have to exactly match the impedances of an implanted leadsystem to the diverter circuits of the present invention. As previouslymentioned, implanted leads usually tend to be inductive, although incertain cases they can even be capacitive. What is important is that thediverter circuit has a reactance which is vectorially opposite to thecharacteristic reactance of the implanted lead. In other words, if theimplanted lead is inductive, it will have a +jωL inductive reactance inohms. One would balance this with a −j/ωC capacitive reactance in thediverter circuit. In an ideal case, the reactance of the divertercircuit would be generally equal and opposite to the characteristicreactance of the implanted lead. In an absolutely ideal situation, theimplanted lead would have a characteristic inductive reactance and thediverter circuit would have an equal but opposite vector quantitycapacitive reactance which would cancel. In order to obtain optimalenergy transfer to an EDS surface in this case, it would further enhanceenergy transfer if the diverter circuit also had a resistive value thatis equal to the characteristic resistance of the implanted lead.Fortunately, when used in combination with a handstop filter, it is notessential that the impedance or reactance of the diversion circuit becompletely equal and opposite to the impedance or reactance of theimplanted lead system.

The present invention is ideal for claiming MRI compatibility for arange of implanted leads. Using a cardiac pacemaker as an example, onemay either through measurement or modeling characterize the impedance ofleads of various lengths, such as 35 to 55 centimeters, and also analyzetheir characteristic impedance over various implant anatomicalgeometries. One could then determine an average impedance or reactanceof this range of leads in order to design an averaged or optimizeddiverter circuit. Unlike for bandstop filters, the diverter circuit willgenerally work over a broad range of circuits, not just a singlefrequency. Accordingly, by using a properly tuned diverter circuitcoupled to an energy dissipation surface of the present invention, onewould be able to assure that a range of lead lengths, lead types andimplant geometries will all be safe in a high electric magnetic fieldenvironment such as MRI.

In a first order approximation, the diverter circuit of the presentinvention can simply be a resistor which is attached to thecharacteristic resistance of the average of the implanted leads forwhich one claims compliance. For example, if the implanted leadsgenerally have a resistance value of around 80 ohms, then one couldachieve a very high degree of tuned energy balance with the presentinvention by having an 80 ohm resistor be coupled between the lead andthe energy dissipating surface. This would not cancel the reactance ofthe lead system but would still go a long way to remove energy from theleads and transfer it to the EDS surface.

A way around this is to use what is known in the industry as an L-Cseries trap filter in combination with the prior art feedthroughcapacitor. When an inductor and a capacitor appear in series, it willalways be a single frequency at which the inductive reactance is equaland opposite to the capacitive reactance. At this point, the series L-Ctrap filter is said to be in resonance. For an ideal series L-C trapfilter (one containing zero resistance), at resonance, it would presenta short circuit. U.S. Pat. No. 6,424,234 describes L-C trap filters(also known as notch filters). The '234 patent describes notch filtersfor a completely different purpose and application. FIG. 10 of U.S. Pat.No. 6,424,234 shows notch filter attenuation in the kilohertz frequencyrange. The reason for this was to provide some degree of attenuationagainst low frequency emitters, such as 58 kHz electronic articlesurveillance (store security) gates. These gates detect tags oncommercial items (such as clothing) as an anti-theft detection system.However, in the present invention, L-C trap filters can be used incombination with a prior art feedthrough capacitor and optimally tunedto dissipate the RF pulsed energy from an MRI system. For example, froma 1.5 Tesla system, the L-C trap filter would be tuned at the Lamourfrequency of 64 MHz. One could also use multiple trap filters within theAIMD or the AIMD header block such that a short circuit or a lowimpedance was provided to multiple MRI systems such as 1.5 Tesla (64MHz); 3 Tesla (128 MHz); 4 Tesla (170 MHz) or 5 Tesla (213 MHz). Allrealizable L-C trap filters have a series resistance. This seriesresistance comes from the resistance of the inductor windings or fromthe equivalent series resistance (ESR) of the capacitor, or both. It isa feature of the present invention, that the resistance of an L-C trapfilter, when used as an energy tuning element to an EDS surface,approximate the characteristic resistance of the implanted lead. Inaccordance with Thevenin's maximum power transfer theorem, this willdissipate maximum energy at the selected MRI frequency to the EESsurface.

The present invention includes switched frequency selective diversion(decoupling) circuits which transfer RF energy which is induced ontoimplanted leads from a high power electromagnetic field environment suchas an MRI RF field to an energy dissipating surface (EDS). In this way,RF energy can be shunted harmlessly into an EDS surface, the AIMDhousing or the bulk thermal mass or handle of a probe or catheter. Thus,the RF or thermal energy can be dissipated in muscle tissue or bodytissues distant from the distal electrodes, or even into flowing bloodor other body fluids, thereby directing such energy away from animplanted lead and especially its tissue contact electrodes. Thediversion (decoupling) circuits of the present invention may also becombined with impeding circuits which can raise and further control theoverall impedance of the system to achieve maximal energy transfer andminimum thermal rise in the implanted lead system.

In other words, an energy dissipating surface is provided with means fordecoupling RF signals from implantable leadwires selectively to saidenergy dissipating surface. In previous studies, concerns have beenraised about the safety of using metallic structures in MR scanners.Radio frequency energy (MHz), transmitted from the scanner in order togenerate the MR signal, can be deposited on the interventional device.This results in high electrical fields around the instrument and localtissue heating. This heating tends to be most concentrated at the endsof the electrical structure. This is certainly true of the implantedleadwires associated with AIMDs. We can address this safety issue usingthe tuned energy balance methods and switched diverter circuits of theinvention. The concern is that the lead electrodes, which directlycontact the tissue, could cause local tissue changes including burns.The present invention is extended beyond the leadwires of probes andcatheters to include the distal tip electrodes associated with theimplanted leads of devices such as pacemakers, cardioverterdefibrillators, neurostimulators and the like. All of these devices havea distal electrode which contacts body tissue in order to deliver pacingpulses or sense biologic activity. It is extremely important that thatinterface junction not overheat and cause localized tissue damage orburning.

U.S. 2003/0050557 explains the need to cut/remove the electrodes fromthe circuit in the MHz frequency range. This is accomplished with theinductor circuit elements. In the MHz frequency range, the surface ringelectrodes are not connected to the rest of the electrical leads.Therefore, the ends of the leads are now buried inside of the catheter.The coupled high electric fields will now be located inside of thecatheter instead of in the tissue. This results in significant reductionand unwanted tissue heating.

In U.S. 2003/0050557, the inside of the catheter, of course, includes abody with a specific thermal mass and specific thermal properties. Overtime, it will rise in temperature and therefore heat surrounding bodytissue. However, this temperature rise is minimal, due to the large areaand thermal mass of the catheter which acts as an energy dissipatingarea or surface. Also, any such minimal heating that does occur is inbody tissue in an area that is distant from the therapy electrode(s).Therefore, the ability for the pacing or stimulus electrode to deliveryenergy in the proper location will not be compromised. By spreading theRF energy over a larger energy dissipating surface area (i.e. inside thecatheter or to an AIMD housing) the temperature rise is thereforereduced and the resulting small amount of heat is generally dissipatedinto bulk body tissues instead of at a specific point.

This is accomplished through switched energy diverting circuits such asbroad band filtering such as capacitive low pass filters, or by resonantfiltering such as creating resonant diverter (trap) circuits consistingof a series inductor and capacitor (L-C trap). Diverting circuits workbest with electrode protecting bandstop filters as described in U.S.Pat. No. 7,363,090; US 2007/0112398 A1; US 2008/0071313 A1; US2008/0049376 A1; US 2008/0132987 A1; and US 2008/0116997 A1, thecontents of which are incorporated herein by reference.

There are three types of electromagnetic fields used in an MRI unit. Thefirst type is the main static magnetic field designated B.sub.0 which isused to align protons in body tissue. The field strength varies from 0.5to 3 Tesla in most of the currently available MRI units in presentclinical use. The second electromagnetic field is the pulsed RF fieldwhich is given by the Lamor Frequency. The Lamor Frequency formula forhydrogen is 42.56 (static field strength in Tesla)=RF frequency. Forexample, for a 1.5 Tesla common hydrogen (proton) scanner, the frequencyof the pulsed RF field is approximately 64 MHz. The third type of fieldis the gradient field which is used to control where the slice is thatgenerates the image that is located within body tissue.

The present invention is primarily directed to the pulsed RF fieldalthough it also has applicability to the gradient field as well.Because of the presence of the powerful static field, non-ferromagneticcomponents are used throughout the present invention. The use offerromagnetic components is contraindicative because they have atendency to saturate or change properties in the presence of the mainstatic field.

FIG. 1 illustrates various types of active implantable medical devicesreferred to generally by the reference numeral 100 that are currently inuse. FIG. 1 is a wire formed diagram of a generic human body showing anumber of exemplary implanted medical devices. 100A is a family ofimplantable hearing devices which can include the group of cochlearimplants, piezoelectric sound bridge transducers and the like. 100Bincludes an entire variety of neurostimulators and brain stimulators.Neurostimulators are used to stimulate the Vagus nerve, for example, totreat epilepsy, obesity and depression. Brain stimulators are similar toa pacemaker-like device and include electrodes implanted deep into thebrain for sensing the onset of the seizure and also providing electricalstimulation to brain tissue to prevent the seizure from actuallyhappening. 100C shows a cardiac pacemaker which is well-known in theart. 100D includes the family of left ventricular assist devices(LVAD's), and artificial hearts, including the recently introducedartificial heart known as the Abiocor. 100E includes an entire family ofdrug pumps which can be used for dispensing of insulin, chemotherapydrugs, pain medications and the like. Insulin pumps are evolving frompassive devices to ones that have sensors and closed loop systems. Thatis, real time monitoring of blood sugar levels will occur. These devicestend to be more sensitive to EMI than passive pumps that have no sensecircuitry or externally implanted leadwires. 100F includes a variety ofimplantable bone growth stimulators for rapid healing of fractures. 100Gincludes urinary incontinence devices. 100H includes the family of painrelief spinal cord stimulators and anti-tremor stimulators. 100H alsoincludes an entire family of other types of neurostimulators used toblock pain. 1001 includes a family of implantable cardioverterdefibrillator (ICD) devices and also includes the family of congestiveheart failure devices (CHF). This is also known in the art as cardioresynchronization therapy devices, otherwise known as CRT devices. 100Jillustrates an externally worn pack. This pack could be an externalinsulin pump, an external drug pump, an external neurostimulator or evena ventricular assist device. 100K illustrates an entire family ofprobes, catheters, venous insert devices such as femoral ICDs, ablationcatheters, loop recorders, and the like.

Referring to US 2003/0050557, Paragraphs 79 through 82, the contents ofwhich are incorporated herein, metallic structures, particularlyleadwires, are described that when placed in MRI scanners, can pick uphigh electrical fields which results in local tissue heating. Thisheating tends to be most concentrated at the ends of the electricalstructure (either at the proximal or distal lead ends). This safetyissue can be addressed using the disclosed systems and methods of thepresent invention. The concern is that the distal electrodes, whichdirectly contact body tissue, can cause local tissue burns. FIGS. 1Athrough 1G in U.S. 2003/0050557 have been redrawn herein as FIGS. 2through 11 and are described as follows in light of the presentinvention.

As used herein, the lead means an implanted lead, including itselectrodes that are in contact with body tissue. In general, for anAIMD, the term lead means the lead that is outside of the AIMD housingand is implanted or directed into body tissues. The term leadwire asused herein, refers to the wiring that is generally inside of an AIMDgenerally between its hermetic terminal and circuit board substrates orinternal circuitry. The term leadwire can also be inclusive of leadwiresinside of a probe or catheter body or handle.

FIG. 2 is a diagrammatic view of a typical prior art device 102 such asa probe, catheter or AIMD lead distal electrode. There are two leadwires104 and 106 which thread through the center of the illustrative probe,catheter or AIMD lead and terminate respectively in a corresponding pairof distal conductive electrode rings 108 and 110. Leadwires 104 and 106are electrically insulated from each other and also electricallyinsulated from any metallic structures located within the catheter orlead body. The overall catheter or implanted lead body is generallyflexible and is made of biocompatible materials, which also havespecific thermal properties. In addition to flexibility, probes andcatheters are typically steerable. AIMD implanted leads are generallymore flexible and are implanted by first placing guide wires. It is wellknown that a push-pull wire (not shown in FIG. 2) can be run down thecenter of the catheter or probe in a lumen and then be attached to acatheter handle or pistol grip or other device so that the physician cancarefully steer or thread the probe or catheter through the torturouspath of the venous system, even into the ventricles of the heart. Suchprobes and catheters, for example, can be used for electrical mappinginside of a heart chamber, or for application of RE energy for ablation,which is used to treat certain cardiac arrhythmias. Probes and cathetershave wide application to a variety of other medical applications. Thereare also combined catheters that can do electrical mapping and can alsoperform RF ablation. When the physician finds the area of arrhythmicelectrical activity and wishes to ablate, he activates a switch whichapplies RF energy to the tip of the catheter (see, e.g., FIG. 55, whichwill be discussed herein in more detail). This would involve a thirdelectrode right at the catheter tip of FIG. 2 (not shown). It would beextremely valuable if the catheter could be guided during real-time MRIimaging. This is important because of MRI's incredible ability to imagesoft tissue. In addition, when one is doing deliberate ablation, forexample, around a pulmonary vein, it is important that a full circle ofscar tissue be formed, for example, to stop atrial fibrillation. MRI hasthe ability to image the scar as it is being formed (for example, seeU.S. Pat. No. 7,155,271). However, it would be highly undesirable if theMRI RE energy that is coupled to the leadwires caused the distalablation tip or the electrode rings to overheat at an improper time,which could burn or ablate healthy tissues.

FIG. 3 shows the interior taken from FIG. 2 showing leadwires 104 and106 which are routed to the two distal electrodes 108 and 110 aspreviously described in FIG. 2.

FIG. 4 shows the electrical circuit of FIG. 3 with a general frequencyselective diverting impedance element 112 connected between leadwires104 and 106. In the present invention, the impedance element 112 canconsist of a number of frequency selective elements as will be furtherdescribed. In general, the first conductive leadwire 104 is electricallycoupled to the first electrode 108, the second conductive leadwire 106is electrically coupled to the second electrode 110, and the frequencydependent diverting reactive element 112 electrically couples the firstand second leadwires 104 and 106 such that high frequency energy isconducted between the first leadwire 104 and the second leadwire 106.

Referring once again to FIG. 4, the frequency selective reactivediverting element 112 tends to be electrically invisible (i.e., a veryhigh impedance) at selected frequencies. The reactive element isdesirably selective such that it would not attenuate, for example, lowfrequency biological signals or RF ablation pulses. However, for highfrequency MRI RF pulsed frequencies (such as 64 MHz), this frequencyreactive element 112 would look more like a short circuit. This wouldhave the effect of sending the energy induced into the leadwires 104 and106 by the MRI RF field back into the catheter body energy dissipatingsurface into which the leadwires are embedded. In other words, there aredesirably both RF energy and thermal conductivity to the probe orcatheter body or sheath or shield which becomes an energy dissipatingsurface all along the lengths of leadwires 104 and 106 such that MRIinduced energy that is present in these leadwires is diverted andconverted to heat into the interior and along the catheter body itself.This prevents the heat build up at the extremely sensitive locationsright at the ring electrodes 108 and 110 which are in intimate anddirect contact with body tissue. In addition, the amount of temperaturerise is very small (just a few degrees) because of the energy beingdissipated over such a relatively high surface area. As previouslymentioned, the high frequency RE pulsed energy from an MRI system cancouple to implanted leads. This creates electromagnetic forces (EMFs)which can result in current flowing through the interface betweenelectrodes that are in contact with body tissue. If this current reachessufficient amplitude, body tissue could be damaged by excessive REcurrent flow or heat build-up. This can create scar tissue formation,tissue damage or even tissue necrosis such to the point where the AIMDcan no longer deliver appropriate therapy. In certain situations, thiscan be life threatening for the patient.

FIG. 5 shows a capacitor 114 which represents one form of the frequencyselective diverting reactive element 112 previously described in FIG. 4.In this case, the reactive element 112 comprises a simple capacitor 114connected between the first conductor or leadwire 104 and the secondconductor or leadwire 106 and will have a variable impedance vs.frequency. The following formula is well known in the art:X_(C)=1/(2πfc). Referring to the foregoing equation, one can see thatsince frequency (f) is in the denominator, as the frequency increases,the capacitive reactance in ohms decreases. With a large number in thedenominator, such as the RE pulsed frequency of a 1.5 Tesla MRI system,which is 64 MHz, the capacitive reactance drops to a very low number(essentially a short circuit). By shorting the leadwires together atthis one frequency, this diverts and prevents the RE energy fromreaching the distal ring electrodes 108 and 110 and being undesirablydissipated as heat into body tissue. Referring once again to FIG. 4, onecan see that the frequency selective diverting element 112 therebydiverts the high frequency RF energy back into the leadwires 104 and106. By spreading this energy along the length of leadwires 104 and 106,it is converted to heat, which is dissipated into the main body of theprobe, catheter or energy dissipating sheath. In this way, therelatively large thermal mass of the probe or catheter becomes an energydissipating surface and any temperature rise is just a few degrees C. Ingeneral, a few degrees of temperature rise is not harmful to bodytissue. In order to cause permanent damage to body tissue, such as anablation scar, it generally requires temperatures above 20° C. Insummary, the frequency selective reactive element 112, which maycomprise a capacitor 114 as shown in FIG. 5, forms a diversion circuitsuch that high frequency energy is diverted away from the distalelectrodes 108 and 110 along the leadwires 104 and 106 to a surface thatis distant from the electrodes 108 and 110, at which point the energy isconverted to heat.

FIG. 6 describes a different way of diverting high frequency energy awayfrom the electrodes 108, 110 and accomplishing the same objective. Thegeneral diverting reactance element 112 described in FIG. 4 is shown inFIG. 6 to comprise a capacitor 114 in series with an inductor 116 toform an L-C trap circuit. For the L-C trap, there is a particularfrequency (f_(r)) at which the capacitive reactance X_(C) and theinductive reactance X, are vectorally equal and opposite and tend tocancel each other out. If there are no losses in such a system, thisresults in a perfect short circuit between leadwires 104 and 106 at theresonant frequency. The frequency of resonance of the trap filter isgiven by the equation f_(r)=

$\frac{1}{2\; \pi \sqrt{LC}},$

wherein f_(r) is the frequency of resonance in Hertz, L is theinductance in henries, and C is the capacitance in farads.

FIG. 7 illustrates any of the aforementioned frequency dependentdiverting impedance elements 112 with the addition of series frequencyselective impeding reactances 118 and 120. The addition of seriesimpedance further impedes or blocks the flow of high frequency MRIinduced currents to the ring electrodes 108 and 110 as will be morefully described in the following drawings.

FIG. 8 is the low frequency model of FIG. 4, 5 or 6. In this regard,FIG. 8 is identical to FIG. 3, in that, once again it shows theelectrical leadwires 104 and 106 connected to the distal ring electrodes108 and 110 of the probe or catheter 102. In the low frequency model,the frequency reactive diverting impedance elements 112 disappearbecause at low frequency their impedances approach infinity. Of course,elongated leads in a probe or catheter are electrically and evenfunctionally equivalent to leads used for cardiac pacemakers,implantable cardioverter defibrillators, neurostimulators and the like.For example, reference is made to U.S. Pat. No. 7,363,090, the contentsof which are incorporated herein. Accordingly, any discussion hereinrelated to probes or catheters apply equally to leadwires for all activeimplantable medical devices as described in FIG. 1, and vice versa.Referring once again to FIG. 8, this is also the low frequency model ofthe circuits shown in FIG. 7. At low frequency, the frequency selectiveor reactive diverting component 112 tends to look like a very high orinfinite impedance. At low frequency, the series reactive or frequencyvariable impeding elements 118 and 120 tend to look like a very lowimpedance or short circuit. Accordingly, they all tend to disappear asshown in FIG. 8.

FIG. 9 is a high frequency model that illustrates how the distalelectrodes or rings 108 and 110 are electrically isolated at highfrequency by shorting leadwires 104 and 106 at location 122. Aspreviously mentioned, such shorting or current diverting could beaccomplished by a capacitor, a capacitive low pass filter or a seriesresonant L-C trap circuit. FIG. 9 also shows the electrodes 108 and 110as cut or disconnected and electrically isolated from the rest of thecircuit. This is because, at very high frequency, series impedingelements 118 and 120 tend to look like a very high impedance or an opencircuit. In summary, by reactive elements 112, 118 and 120 actingcooperatively, reactive element 112 diverts the high frequency energyback into energy dissipating surfaces in the probe or catheter while atthe same time reactive elements 118 and 120 impede the high frequency RFenergy. Accordingly, in the ideal case, at high frequencies, theequivalent circuit of FIG. 9 is achieved. Accordingly, excessive highfrequency MRI RF energy cannot reach the distal ring electrodes 108, 110and cause undesirable heating at that critical tissue interfacelocation.

FIG. 10 shows any of the previously described diverting frequencyselective impedance elements 112 in combination with series reactancecomponents shown in the form of a pair of inductors 116 a, 116 b. It iswell known to electrical engineers that the inductive reactance in ohmsis given by the equation X_(L)=2πfL. In this case the frequency term (f)is in the numerator. Accordingly, as the frequency increases, thereactance (ohms) of the inductors also increases. When the frequency isvery high (such as 64 MHz) then the reactance in ohms becomes extremelyhigh (ideally approaches infinity and cuts off the electrodes). Byhaving a short circuit or very low impedance between the leadwires andthe probe/catheter body 104 and 106 and then, at the same time, having avery high impedance in series with the electrodes from inductors 116,this provides a very high degree of attenuation to MRI RF pulsedfrequencies thereby preventing such energy from reaching the distal ringelectrodes 108 and 110. In FIG. 10, the line-to-line selective impedanceelement 112 diverts high frequency energy back into leadwires 104 and106 while at the same time the series inductors 116 impede (or cut-off)high frequency energy. When the line-to-line element 112 is a capacitor114 as shown in FIG. 5, then this forms what is known in the prior artas an L section low pass filter, wherein the capacitor 114 electricallycooperates with the inductors 116 (FIG. 10) to form a 2-element low passfilter. By definition, a low pass filter allows low frequencies such asbiological signals to pass to (stimulation pulses) and from (biologicsensing) the distal electrodes freely without attenuation while at thesame time providing a high degree of attenuation to undesirable highfrequency energy. It will be obvious to those skilled in the art thatFIG. 5 describes a single element (capacitor) low pass filter, and thatFIG. 10 describes a 2-element or L-section low pass filter. Moreover,any number of inductor and capacitor combinations can be used for lowpass filters, including 3-element Pi or T circuits, LL, 5-element oreven “n” element filters.

FIG. 11 offers an even greater performance improvement over thatpreviously described in FIG. 10. In FIG. 11, modified frequencyselective impeding elements each incorporate a parallel resonantinductor 116 and capacitor 114 which is also known in the industry as abandstop filter 117. The L-C components for each of the reactiveelements are carefully chosen such that each of the bandstop filters 117are resonant, for example, at the pulsed resonant frequency of an MRIscanner. For common hydrogen scanners, the pulsed resonant frequency ofan MR scanner is given by the Lamor equation wherein the RF pulsedfrequency in megahertz is equal to 42.56 times the static fieldstrength. For example, for a popular 1.5 Tesla scanner, the RF pulsedfrequency is 64 MHz. Common MR scanners that are either in use or indevelopment today along with their RF pulsed frequencies include: 0.5Tesla-21 MHz; 1.5 Tesla-64 MHz; 3 Tesla-128 MHz; 4 Tesla-170 MHz; 5Tesla-213 MHz; 7 Tesla-300 MHz; 8 Tesla-340 MHz; and 9.4 Tesla-400 MHz.When the bandstop filters 117 are resonant at any one of these RF pulsedfrequencies, then these elements tend to look like an open circuit whichimpedes the flow of RF current to distal electrodes. When compatibilitywith different types of MR scanners is required, for example, 1.5, 3 and5 Tesla, then three separate bandstop filter elements in series maycomprise the reactive element 118 (FIG. 7), and three separate bandstopfilter elements in series may comprise the reactive element 120 (FIG.7). Each of these would have their L and C components carefully selectedso that they would be resonant at different frequencies. For example, inthe case of MR scanners operating at 1.5, 3 and 5 Tesla, the threebandstop filters comprising the reactive element 118 as well as thethree bandstop filters comprising the reactive element 120 would beresonant respectively at 64 MHz, at 128 MHz, and at 170 MHz. Theresonant frequencies of the bandstop filter elements could also beselected such that they are resonant at the operating frequency of otheremitters that the patient may encounter such as diathermy and the like.The use of bandstop filters 117 is more thoroughly described in U.S.Pat. No. 7,363,090; US 2007/0112398 A1; US 2007/0288058; US 2008/0071313A1; US 2008/0049376 A1; US 2008/0161886 A1; US 2008/0132987 A1 and US2008/0116997 A1, the contents of which are incorporated herein.

Referring now to FIG. 12, a prior art active implantable medical device(AIMD) 100C is illustrated. In general, the AIMD 100C could, forexample, be a cardiac pacemaker which is enclosed by a titanium orstainless steel housing 124 as indicated. The titanium housing 124 ishermetically sealed and contains circuit board 137, however, there is apoint where conductors such as the illustrative conductors 126 a, 126 b,126 c and 126 d must ingress and egress in non-conductive relationshiprelative to the housing 124. This is accomplished by providing ahermetic terminal assembly 128. Hermetic terminal assemblies 128 arewell known and generally consist of a ferrule 129 which is laser weldedto the titanium housing 124 of the AIMD 100C. In FIG. 12, fourconductors 126 a-126 d are shown for connection to a correspondingnumber of leadwires, such as the illustrative bipolar leads 104 and 106shown for coupling to the connector receptacles 130. In thisconfiguration, the four leadwires coupled respectively to the conductors126 a-126 d comprise a typical dual chamber bipolar cardiac pacemaker.It should be noted that each of the bipolar leads 104 and 106 have apair of leadwires associated with them. These are known as bipolarelectrodes wherein one wire is routed to the tip electrode and the otheris routed to the ring electrode in locations 108 and 110.

Connectors 132 are commonly known as IS-1 connectors and are designed toplug into mating receptacles 130 on a header block 134 mounted on thepacemaker housing 124. These are low voltage (pacemaker) leadwireconnectors covered by an International Standards Organization (ISO)standard IS-l. Higher voltage devices, such as implantable cardioverterdefibrillators, are covered by a standard known as the ISO DF 1. A newstandard was recently published that will integrate both high voltageand low voltage connectors into a new miniature quadpolar connectorseries known as the ISO IS-4 standard. Leads plugged into theseconnectors are typically routed in a pacemaker or ICD application downinto the right ventricle and right atrium of the heart 136. There arealso new generation devices that have been introduced to the market thatcouple leads to the outside of the left ventricle. These are known asbiventricular devices and are very effective in cardiacresynchronization therapy (CRT) and treating congestive heart failure(CHF).

It should be obvious to those skilled in the art that all of thedescriptions herein are equally applicable to other types of AIMDs.These include implantable cardioverter defibrillators (ICDs),neurostimulators, including deep brain stimulators, spinal cordstimulators, cochlear implants, incontinence stimulators and the like,and drug pumps. The present invention is also applicable to a widevariety of minimally invasive AIMDs. For example, in certain hospitalcath lab procedures, one can insert an AIND for temporary use such as aprobe, catheter or femoral artery ICD. Ventricular assist devices alsocan fall into this type of category. This list is not meant to belimiting, but is only example of the applications of the noveltechnology currently described herein. In the following description,functionally equivalent elements shown in various embodiments will oftenbe referred to utilizing the same reference number.

FIG. 13 illustrates a prior art single chamber bipolar AIMD 100C andleadwires 104 and 106 with a distal tip electrode 138 and a ringelectrode 108 typically as used with a cardiac pacemaker 100C. Shouldthe patient be exposed to the fields of an MRI scanner or other powerfulemitter used during a medical diagnostic procedure, currents that aredirectly induced in the leadwires 104, 106 can cause heating by I²Rlosses in the leadwires or by heating caused by RF current flowing fromthe tip and ring electrodes 138, 108 into body tissue. If these inducedRF currents become excessive, the associated heating can cause damage oreven destructive ablation to body tissue 136.

FIG. 14 illustrates a single chamber bipolar cardiac pacemaker 100C, andleadwires 104 and 106 having distal tip 138 and distal ring electrode108. This is a spiral wound (coaxial) lead system where the tipelectrode leadwire 104 is wrapped around the ring electrode leadwire106. The characteristic impedance of this lead type usually has aninductive component. There are other types of pacemaker lead systems inwhich these two leadwires that lay parallel to one another (known as abifilar lead system), which are not shown.

FIG. 15 is an enlarged schematic illustration of the area “15-15” inFIG. 14. In the area of the distal tip 138 and ring electrode 108,bandstop filters 117 a, 117 b have been placed in series with each ofthe respective ring and tip circuits. The ring circuit leadwire 104 hasbeen drawn straight instead of coiled for simplicity. The bandstopfilters 117 are tuned such that, at an MRI pulsed RF frequency, a highimpedance will be presented thereby reducing or stopping the flow ofundesirable MRI induced RF current from the electrodes 138 and 108 intobody tissues.

The tip electrode 138 is designed to be inserted into intimate contactwith myocardial tissue. Over time it can become encapsulated and fullyembedded or buried within such tissue. However, the ring electrode 108is designed to float within the blood pool, for example, in a cardiacchamber such as a ventricle or atrium. With the constant bloodperfusion, the ring electrode 108 can be somewhat cooled during medicaldiagnostic procedures, such as MRI. However, the tip electrode 138,which is embedded in the myocardial tissue, is thermally insulated incomparison. Moreover, it can't always be assumed that a ring electrode108 that is floating in the blood pool will be adequately cooled by theflow of blood. There are certain types of patients that havecardiovascular diseases that lead to very low election fractions and lowblood flow rates and even perfusion issues. The ring electrode 108 canalso become encapsulated by body tissues. Accordingly, both the distaltip electrode 138 and the ring electrode 108 are preferably bothassociated with bandstop filters 117 a, 117 b. However, since theoperation of the bandstop filter 117 is more important with the tipelectrode 138 than it is with the ring electrode 108, in order toprevent distal tip heating and associated tissue damage, in many AIMDapplications only a tip bandstop filter 117 a may be required for MRIcompatibility.

FIG. 16 is a front and side view tracing of an actual patient X-ray.This particular patient required a cardiac pacemaker. The correspondingimplantable leadwire system, as one can see, makes for a verycomplicated antenna and loop coupling situation. The reader is referredto the article entitled, “Estimation of Effective Lead Loop Area forImplantable Pulse Generator and Implantable Cardioverter Defibrillators”provided by the AAMI Pacemaker EMC Task Force. In FIG. 16, one can seefrom the X-ray tracing that there are electrodes in both the rightatrium and in the right ventricle. Both these involve a separate tip andring electrode (not shown in FIG. 16). In the industry, this is known asa dual chamber bipolar leadwire system. It will be obvious to thoseskilled in the art that any of the passive frequency selective networks,as previously described in FIGS. 2 through 11, can be incorporated intothe leadwires as illustrated in the X-ray tracing of FIG. 16. Frequencyselective diverter and/or impeding filters of FIGS. 2-11 of the presentinvention are needed so that MRI exposure cannot induce excessivecurrents into the associated leads or electrodes. There are also newercombined pacemaker/ICD systems which include biventricular pacemaking(pacing of the left ventricle). These systems can have as many as 12implanted leads, 140.

FIG. 17 is a line drawing of an actual patient cardiac X-rav of one ofthe newer bi-ventricular leadwire systems with various types ofelectrode tips 140 shown. For instance, electrode tip 140 a is a passivefixation right atrium pacing lead, electrode tip 140 b is an activefixation bi-ventricular pacing lead, and electrode tip 140 c is aventricle defibrillation lead. The new bi-ventricular systems are beingused to treat congestive heart failure, and make it possible to implantleads outside of the left ventricle. This makes for a very efficientpacing system; however, the implantable leadwire system is quitecomplex. When a leadwire system, such as those described in FIGS. 12-17,are exposed to a time varying electromagnetic field, electric currentscan be induced into the electrodes of such leadwire systems. For thebi-ventricular system, a passive component frequency diverting networkof FIGS. 2-11 would need to be placed in conjunction with each of thethree distal tips and ring electrodes to corresponding energydissipating surfaces.

The word passive is very important in this context. Active electroniccircuits, which are defined as those that require power, do not operatevery well under very high amplitude electromagnetic field conditions.Active electronic filters, which generally are made from microelectronicchips, have very low dynamic range. Extremely high fields inside an MRIchamber would tend to saturate such filters and make them becomenonlinear and ineffective. Accordingly, frequency selective networks arepreferably realized using non-ferromagnetic passive component elements.In general, this means that the frequency selective components for bothdiverters and impeders preferably consist of capacitors, inductors, andresistors in various combinations. Passive component elements arecapable of handling very high power levels without changing theircharacteristics or saturating. Moreover, the inductor elements arepreferably made from materials that are not ferromagnetic. The reasonfor this is that MRI machines have a very powerful main static magneticfield (B₀). This powerful static magnetic field tends to saturateferrite elements and would thereby change dramatically the value of theinductance component. Accordingly, in the present invention, theinductor elements are preferably fabricated without the use of ferrites,nickel, iron, cobalt or other similar ferromagnetic materials that arecommonly used in general electronic circuit applications.

FIG. 18 gives the frequency of resonance f_(r) for the parallel L-Cbandstop filter circuit 117 of FIG. 15: where f_(r) is the frequency ofresonance in Hertz, L is the inductance in Henries and C is thecapacitance in Farads. The same equation given in FIG. 18 also appliesto a series L-C trap filter illustrated as a frequency diverter element112 in FIG. 6. The inductor L is designated by 116 and the capacitor Cis designated by 114 in FIG. 6. Clinical MRI systems vary in staticfield strength from 0.5 Tesla all the way up to 3 Tesla with newerresearch machines going as high as 11.4 T. The frequency of the pulsedRF field associated with the static field is given by the LamourEquation, f=γT, where T is the field strength in Teslas, and γ isgyromagnetic ratio for hydrogen, which is 42.58 MHz/T. Accordingly, a 3Tesla MRI system has a pulsed RF field of approximately 128 MHz.

By referring to FIG. 18, one can see that the resonant frequency fr ofan ideal tank filter can be predicted by using the equation:

$\frac{1}{2\; \pi \sqrt{LC}}$

Where f_(r) is the resonant frequency, L is the inductance, in Henries,of the inductor component, and C is the capacitance, in Farads, of thecapacitor component. In this equation, there are three variables: f_(r),L, and C. The resonant frequency, f_(r), is a function of the MRI systemof interest. As previously discussed, a 1.5 T MRI system utilizes an RFsystem operating at approximately 64 MHz, a 3.0 T system utilizes a 128MHz RF, and so on. By determining the MRI system of interest, only L andC remain. By artificially setting one of these parameters, a filterdesigner needs only to solve for the remaining variable.

FIG. 19 is a graph showing the impedance versus frequency for a bandstopfilter as previously described in U.S. Pat. No. 7,363,090 and US2007/0112398 A1. One can see that at frequencies outside of the resonantfrequency, impedance is nearly zero ohms. When the capacitor andparallel inductor is in resonance, the impedance is very high (ideallyinfinity).

FIG. 20 is the impedance equation for the capacitor in parallel with aninductor of a bandstop filter.

FIG. 21 gives the equations for inductive reactance X_(L) and capacitivereactance X₀. In all of these equations (f) is frequency in hertz, L isinductance in henries, and C is capacitance in farads.

It should also be noted that the L-C parallel bandstop filter alsocaptures the RF energy. That is, for example, for a 1.5 Tesla system,the energy will swap back and forth between the capacitor and theinductor at 64 MHz The energy is stored first in the capacitor's staticfield and then discharged into the magnetic field of the inductor andback and forth. Accordingly, relatively high currents can circulate backand forth between these two circuit elements during an MRI procedure.Accordingly, it is important that these two components be robust enoughto handle these currents. It is also important that their resistiveelements be relatively low such that the bandstop filter itself does notbecome a heating element within the lead system. One way to increase thecurrent handling capabilities and reduce the resistance of the capacitorelements is the use of dual electrode plates as described in U.S. Pat.No. 5,978,204, the contents of which are incorporated herein. It shouldalso be noted that RF power handling is a special concern for all of thefrequency diverter circuits 112 as illustrated in FIGS. 4, 5, 7, 10, and11. The diverter elements have to be designed in a very robust mannersince they are carrying high levels of RF current to the energydissipating surface EDS. The impeding elements tend to reduce currentand therefore usually are not required to carry extremely high levels ofRF current. For example, referring once again to FIG. 5, once can seethat in this case, the diverter element is a capacitor 114. Thecapacitive reactance is a very low impedance at high frequency, such as64 megahertz wherein relatively high amplitude RF currents flow in andout of the capacitor's internal electrodes. If a capacitor has highequivalent series resistance, or high ohmic loss, the capacitor itselfcould get very hot. One of the ways to reduce the internal resistanceand/or RF current handling capability of the passive diverter elementsof the present invention will be described in detail further on.

FIG. 19 is a graph showing impedance versus frequency for the idealparallel tank circuit 117 of FIG. 15. As one can see, using ideal (zeroresistance) circuit components, the impedance measured between points Aand B for the parallel tank circuit 117 shown in FIG. 15 is zero untilone approaches the resonant frequency f_(r). At the frequency ofresonance, these ideal components combine together to approach aninfinite impedance. This comes from the equation Z_(ab) for theimpedance for the inductor in parallel with the capacitor shown as FIG.20. When the inductive reactance is equal to the capacitive reactance,the two imaginary vectors cancel each other and sum to zero causing theequation to become discontinuous at this point. Referring to theequations in FIGS. 20 and 21, one can see in the impedance equation forZ_(ab), that a zero will appear in the denominator whenX_(L)=X_(c)(jωL=−j/ωC). This has the effect of making the impedanceapproach infinity as the denominator approaches zero. This means that atone unique frequency, the impedance between points A and B in FIG. 19will appear very high (analogous to opening a switch). Accordingly, itwould be possible, for example, in the case of a cardiac pacemaker, todesign the cardiac pacemaker for compatibility with one single popularMRI system. For example, in the patient literature, the device manualand perhaps contained in the digitally stored information on animplanted RFID chip, it could be noted that the pacemaker leadwiresystem has been designed to be compatible with 3 Tesla MRI systems.Accordingly, with this particular device, a distal tip bandstop filter117 would be incorporated where the L and the C values have beencarefully selected to be resonant at 128 MHz, presenting a high oralmost infinite impedance at the MRI pulse frequency.

FIG. 22 is generally taken from FIG. 15 showing a typical prior artbipolar pacemaker leadwire system. Shown is the distal tip electrode 138and ring electrode 108. An insulation or insulative lead body 142 isalso illustrated. The distal tip electrode 138 can be passive (meaningthat it can be bent back in a “J” or shoved against myocardial tissue sothat it just rests against the tissue). A more commonly used electrodetoday is known as the active fixation tip. This is an electrode where byturning the entire center of the lead, the physicians can screw a helixinto myocardial tissue thereby firmly affixing it. A prior art activefixation electrode tip 144 is shown in FIG. 23. This is typically usedin conjunction with a cardiac pacemaker, an implantable defibrillator orthe like. One can see that an active fixation tip housing 146 is pressedup against the tissue to be stimulated, e.g., the myocardial tissue 46of the patient's heart. For example, this could be the septal wallbetween the right ventricle and the left ventricle. A helix electrodeassembly 148 is shown in a retracted position relative to the adjacentheart tissue 46. At the lead proximal end in the pectoral pocket, thephysician uses a tool to axially twist the assembly shaft 150, whichdrives the pointed tip helix screw 152 into the myocardial tissue,firmly affixing it. As can be seen, it would be highly undesirable forthe active fixation helix screw 152 to heat up during an MRI scan.Because the helix screw 152 is deeply embedded into myocardial tissue,if excessive heating and temperature rise did occur, not only couldscarring or ablation of cardiac tissue occur, but an actual cardiac wallperforation or lesion could result in sudden death. It will also beobvious to those skilled in the art that any of the frequency impedingor diverting circuits, as shown in FIG. 4, 5, 6, 7, 10 or 11, would behighly undesirable if they were located within the overall housing 146of the active fixation tip 144. This is because the heat would indeed beremoved from the helix screw 152, but it would be transferred into theactive fixation housing 146 which also rests in intimate contact withthe endocardium heart tissue. What this means is that redirecting theMRI induced electromagnetic energy from the helix tip 152 to the housing146 simply moves the heat from one bad location to another bad location.Because the housing 146 is also in intimate contact with heart tissue,one would experience excessive temperature rise and resulting tissueburning, scarring or necrosis at that location as well.

Referring once again to FIG. 22, one can see that there is a ringelectrode 108 which is placed back (spaced proximally) a suitabledistance from the distal tip electrode 138. In a bipolar pacing system,the cardiac pacing pulse is produced between the tip electrode 138 andthe ring electrode 108. This electrical pulse induced into myocardialtissue produces a heartbeat. Sensing can also be accomplished betweenthese two electrodes 138, 108 wherein the pacemaker can constantlymonitor the electrical activity of the heart. There are similaranalogies for neurostimulators, cochlear implants and the like. There isusually a point at which the distal electrodes, for example electrode138, contact body tissue or fluid for delivery of therapy involvingelectrical energy. In a neurostimulator application, such as a spinalcord stimulator, small electrical currents or pulses are used to blockpain in the spinal nerve roots (create paresthesia). In a urinaryincontinence stimulator, a distal electrode is used to cause a musclecontraction and thereby prevent undesirable leakage of urine. In all ofthese examples, it would be highly undesirable for excess heatingdefined as temperature rise above a few degrees C., to occurparticularly at the implanted lead electrode(s).

In previous studies, concerns have been raised about the safety of usingmetallic structures, such as leadwires and MR scanners. Radio frequencyenergy (MHz) transmitted from the MRI scanner in order to generate theMR signal can be coupled onto the interventional device or itsassociated leads. This results in high electrical fields around theinstrument and local tissue heating. This heating tends to be mostconcentrated at the ends of the electrical structure or leads.

We can address this safety issue using the methods of the presentinvention. The concern is that distal electrodes or distal surface ringelectrodes, which directly contact body tissue, will cause local tissueburns. We need to re-direct the RF induced energy from the leads to anEDS surface. In the current embodiment, this is accomplished primarilywith tuned frequency selective diverter circuit elements to an EDSsurface or housing.

A very effective way to “cut” or impede RF energy current flow toimplanted lead distal electrodes is to use a parallel resonant bandstopfilter circuit in place of the inductors in FIG. 10. This resonantcircuit could consist of an inductor in parallel with a capacitor (anL-C bandstop filter as shown in FIG. 11). If this parallel L-C circuitis tuned to the MR frequency, it will present a very high impedance atthis frequency. This will effectively cut or disconnect the electrodesfrom the elongated leads at the MRI frequency and prevent unwantedheating. For maximal effectiveness, the L-C circuit should be shielded.For a probe or a catheter application, with these design concepts, theelectrical end of the leads (in the MHz range) are buried inside of thecatheter body and as a result, the concentrated electric fields are alsolocated inside of the capacitor, instead of in the tissue. This resultsin a significant reduction in unwanted tissue heating. As previouslymentioned, a resonant circuit is an effective way to “cut” the surfaceelectrodes from the rest of the electrical circuit. This resonantcircuit could be an inductor in parallel with the capacitor (a bandstopfilter also known as an L-C “tank” circuit). Probes and catheters oftenincorporate metallic sheaths which also assist in dissipating theunwanted energy over large surface areas. This is equivalent to theenergy dissipating surface (EDS) structures as described herein. One ofthe advantages of bandstop filters is that they will allow low frequencypacing pulses and biologic sensing signals to freely pass through. Thisis very important for a lifesaving AIMD such as a cardiac pacemaker.However, there are many neurostimulator applications, for example,spinal cord stimulators that might have eight, sixteen or even 24electrodes. These electrodes may have to be placed in the spinal cordnerve root, which is a very small space adjacent to the spine.Accordingly, it becomes very impractical to place that many bandstopfilters into such a small and torturous location. Similar analogiesexist for multiple deep brain stimulation electrodes which must bephysically very small in size to penetrate through deep brain tissuewithout collateral damage. Accordingly, there is a need for a supplementor an alternative to bandstop filters. An ideal solution is the tunedenergy balance system and energy dissipating surfaces of the presentinvention. While optimally or even sub-optimally tuning the frequencyselective diverter elements to an EDS surface, one can draw theRF-induced energy out of the implanted lead system and thereby dissipateit at an EDS surface at a location away from sensitive body tissues. Aperfect example is a spinal cord stimulator. As mentioned, theelectrodes are placed along the spine in the spinal cord nerveroot/canal. The leads routed from these electrodes are generally routedto an AIMD which is typically implanted either in the buttocks or thelower back. This is an ideal situation for the tuned energy balancesystem and EDS of the present invention. For one thing, in an MRIscanner, the human spine is generally located fairly close to the MRIbore iso-center. At iso-center, the RF electric fields in a scanner tendto be quite low in amplitude (nearly zero). Therefore, the inducedenergy on the electrodes is relatively small compared to the RF-inducedenergy along the rest of the lead path. In fact, in this scenario, thehighest electric fields will be furthest from iso-center, which meansthe leads routed into the buttocks or lower back that are adjacent tothe AIMD itself. It will be obvious to those skilled in the art, that itis far preferable to have a slight temperature rise over the relativelylarge surface area (housing) of the AIMD in the buttocks area. This isfar less dangerous to the patient than a temperature rise at theelectrodes that are placed immediately adjacent the spinal cord nerve.Thermal injury to the spinal nerve can cause very serious and lastingneurologic deficits.

All of the circuit elements as described in connection with FIGS. 4through 11 are for purposes of redirecting high frequency RF energy awayfrom lead electrodes into a location that has larger thermal mass andlarger area such that the energy is not being dissipated at theconcentrated point of electrode to tissue contact. Concentrating the MRIRF energy at an electrode causes excessive temperature rise which canresult in damage to adjacent body tissues. Referring back to FIG. 3, onecan see that the leadwires 104 and 106 are embedded in the insulatingsheath of a probe, a catheter, a cardiac pacemaker lead or the like.Accordingly, if excess heat is dissipated along these leadwires, it isthen dissipated into these surrounding structures. As previouslymentioned, there is also a parasitic capacitance that's formed alongthese leadwires and the surrounding structures or insulating sheaths. Itis a feature of the present invention that any of the passive componentfrequency selective circuits can also be directly connected to energydissipating elements that are proximal from the electrodes themselves.

Referring to FIG. 22 (and also FIGS. 24-26), the insulation sheath 142typically encapsulates the leadwires 104 and 106 in silicone orpolyurethane to provide strength and resistance to body fluids. Theinsulation sheath 142 has thermal conduction properties and alsoprovides important electrical isolation between the leadwires 104 and106 themselves and also surrounding body fluids and tissues.

FIG. 24 is generally taken from FIG. 22 except that the twisted orcoaxial lead wires 104 and 106 have been straightened out for betterillustration of the examples of the present invention. This is alsoanalogous to FIG. 2 for the wires of probes and catheters previouslydescribed herein. The straightened and elongated leadwires 104, 106 ofFIG. 24 are also illustrative of certain bifilar leadwire systems, whichcan also be used for pacemakers, neurostimulators and the like. In otherwords, the leadwires are not always twisted as shown in FIG. 22 as thereare certain applications where it is desirable to have the leadwires104, 106 running parallel to each other in a straight fashion. Forillustrative purposes, we will focus on the straight leadwires 104, 106of FIG. 24, but realize that all of these same principles to follow areequally applicable to twisted or coaxial leadwires. In FIG. 22, one cansee that the insulation sheath 142 generally runs up to and fixates thering electrode 108, but does not cover or encapsulate it. This is alsotrue for the distal tip electrode 138. This is important such that theelectrodes are not insulated, so that they can contact body tissue anddeliver therapy and/or sense biologic signals. If they were insulated,they would not be able to function and conduct electrical current intobody tissue. In practice, the parasitic capacitance value is quite low.For differential mode induced EMFs, by electrically shorting leadwires104 and 106 together, the energy induced from an MRI system is containedinto a loop whereby it will create relatively high RF currents inleadwires 104 and 106. Importantly, this loop disconnects this currentflow from the distal electrodes 138 and 108. Accordingly, this energywill be converted to heat within leadwires 104 and 106 where it will bethermally conducted into the insulation sheath 142 and dissipated over amuch larger surface area. In the case where the induced EMFs are commonmode, frequency selective networks diverting of the present inventionare used to couple the high frequency energy to a metallic surface ofthe probe or catheter, such as a shield, or to an equivalent energydissipating surface (EDS). This has the effect of preventing a largetemperature rise at the electrode to tissue interface which could bedamaging to body tissue. More importantly, said RF energy or heat isdiverted away from the distal electrodes, which make direct contact withsensitive body tissues. It is in this location where excessive heatdissipation can cause temperature rises that can cause damage to bodytissue and therefore, undesirable loss of therapy or evenlife-threatening tissue damage. In a preferred embodiment, the parasiticcapacitances or heat conductive interface would be replaced by passivecomponent capacitances that are connected directly to a conductiveenergy dissipating surface. This is a more efficient way of divertingthe energy to a point distant from the distal electrodes and convertingit to heat. By re-directing the RF and/or thermal energy to a point oran area distant from the distal electrodes, one thereby provides a highdegree of protection to the sensitive junction between the electrodesand body tissue. For example, that junction may be the point where adistal electrode contacts myocardial tissue and provides criticallyimportant pacing pulses. Energy concentration at distal electrode cancause dangerous temperature rises.

FIG. 25 is generally equivalent and incorporates and embodies theconcepts previously described in FIGS. 2 through 11 herein. In FIG. 25,one can see the lead insulation 142. There are parasitic capacitances114 which are formed between leadwires 104 and 106 and the insulationlayer 142. At high frequency, this has the desired effect of divertingor shunting high frequency MRI RF energy away from the leadwires 104 and106 thereby redirecting energy into the insulation sheath 142 where itcan be dissipated over a much larger surface area with minimaltemperature rise. Series reactive impeding elements 118 and 120, aspreviously described and shown in connection with FIG. 7, block, cut orimpede the flow of MRI induced RF energy to the distal tip electrode 138and/or the distal ring electrode 108, wherein these electrodes 138, 108correspond respectively with the ring electrodes 108, 110 shown in FIGS.2-11. These series frequency selective reactances 118 and 120 areoptional, but do increase the efficacy of the present system.

Reactance 112 can be a simple capacitor as shown in FIG. 5, a low-passfilter or it can be an L-C series trap filter as shown in FIG. 6. Thistends to short leadwires 104 and 106 together at high frequency therebydiverting undesirable high frequency energy and thereby preventing itfrom reaching distal tip electrode 138 or ring electrode 108. Referringonce again to FIG. 25, we can see high frequency RF currents I andThese, for example, are the RF pulsed currents induced in an elongatedimplanted lead from a 1.5 Tesla MRI system, and they would oscillateback and forth at 64 MHz thereby reversing directions, as shown, at thatfrequency. This is better understood by referring to FIG. 9. Thecurrents are cut off (as indicated at 122 in FIG. 9) and are effectivelycontained within leadwires 104 and 106. This redirects the energy thatis induced by the high frequency MR fields back into the EDS sheath 142at a point distant from the distal electrodes 138 and 108. This EDSdesirably prevents the distal electrodes from overheating at their pointof contact with body tissue.

FIG. 26 is very similar to the structures shown in FIGS. 22 and 24 foractive implantable medical devices (AIMDs) such as cardiac pacemakersand the like. Shown is a frequency selective diverter element 112 inaccordance with FIG. 6, which in this case consists of an inductor 116in series with a capacitor 114 (L-C trap filter). The component valuesof the inductor 116 and the capacitor 114 can be selected such that theyare resonant at a particular frequency. In this case, for illustrativepurposes, they shall be resonant at 64 MHz thereby providing a lowimpedance short circuit for 1.5 Tesla MRI signals. This has the effectof diverting or shunting the energy off of leadwire 104 to therelatively large surface area of the ring electrode 108. The ringelectrode 108 is typically a metallic structure consisting of acylindrical ring and very high thermal conductivity. It also has, bynecessity, very high electrical conductivity. Accordingly, referringonce again to FIG. 26, the ring electrode 108, by its inherent nature,becomes an energy dissipating surface 161 wherein the high frequency RFenergy is diverted to it, wherein said RF energy will either beconverted to heat, which will be directed into the surrounding bloodflow, or said RF energy will be harmlessly dissipated into surroundingbody tissues. More specifically, for example, in the right ventricle,the distal tip electrode 138, 152 is designed to be screwed intomyocardial tissue in accordance with FIG. 23. The ring electrode 108, onthe other hand, is designed to be placed back away from distal tipelectrode 138, 152 such that it actually floats in the pool of bloodthat is flowing in the particular cardiac chamber. In an idealsituation, the wash of blood over it tends to provide a constant coolingaction through heat transfer over the ring electrode 108 therebydissipating undesirable heat from high frequency RF energy harmlesslyinto the flowing blood (or other body fluids such as lymph in otherapplications). A disadvantage of this approach is that in a certainpercentage of patients both the tip and the ring tend to be overgrown bytissue. Accordingly, the use of a separate energy dissipating surface161, which is located further back from both the distal tip and ringelectrode, is desirable such that it is guaranteed to remain in theblood pool. For the energy dissipating surface 161, which can either bethe ring electrode itself or a separate energy dissipating structure161, it is a desirable feature that it includes some type of biomimeticcoating such that tissue overgrowth is inhibited. Referring back to FIG.25, for example, a biomimetic coating 154 could be deposited all overthe ring electrode 108 to thereby inhibit tissue overgrowth.

FIG. 27 is a line drawing of a human heart with cardiac pacemaker dualchamber bipolar leads shown in the right ventricle 156 and the rightatrium 158 of a human heart 136. FIG. 27 is taken from slide number 3from a PowerPoint presentation given at The 28^(th) Annual ScientificSessions of the Heart. Rhythm Society by Dr. Bruce L. Wilkoff, M. D. ofthe Cleveland Clinic Foundation. This article was given in Session 113on Friday, May 11, 2007 and was entitled, ICD LEAD EXTRACTION OFINFECTED AND/OR REDUNDANT LEADS. These slides are incorporated herein byreference and will be referred to again simply as the Wilkoff reference.In FIG. 27, one can see multiple leadwires extending from an activeimplantable medical device 100C (such as a cardiac pacemaker or thelike) coupled to associated electrodes, one of which comprises thedistal tip ventricular electrode 138 located in the right ventricular156 apex. The dark shaded areas in FIG. 25 show the experience of theCleveland Clinic and Dr. Wilkoff (who is a specialist in leadextraction), where extreme tissue overgrowth and vegetation tends tooccur. There are numerous cases of extracted leads where both the tipand ring electrodes have been overgrown and encapsulated by tissue.Referring once again to FIG. 27, one can see tip electrode 138, which islocated in the right ventricular apex. The shaded area encasing thiselectrode 138 shows that this area tends to become encapsulated by bodytissue. A distal tip electrode 144 in the right atrium 158 may similarlybe overgrown and encapsulated by tissue, as shown by the encasing shadedarea. There are other areas in the superior vena cava and venous systemwhere leads tend to be encapsulated by body tissue a great percentage ofthe time. These are shown as areas 157 and 159. This is particularlyimportant to know for the present invention since these would be highlyundesirable areas to place an energy dissipating surface 161 inaccordance with the present invention. Ideal locations for energydissipating surfaces are shown where there tends to be little to notissue overgrowth as 161 a, 161 b, or 161 c.

Referring once again to FIG. 27, as previously mentioned, it is veryimportant that this leadwire system does not overheat during MRIprocedures particularly at or near the distal tip electrodes and rings.If both the distal tip and ring electrode become overgrown by bodytissue, excessive overheating can cause scarring, burning or necrosis ofsaid tissues. This can result in loss of capture (loss pacing pulses)which can be life-threatening for a pacemaker dependent patient. It isalso the case where implanted leads are often abandoned (where the leadhas been permanently disconnected from the AIMD), Often times when thedevice such as a pacemaker 102 shown in FIG. 27 is changed out, forexample, due to low battery life and a new pacemaker is installed, thephysician may decide to install new leadwires at the same time.Leadwires are also abandoned for other reasons, such as a dislodged or ahigh impedance threshold. Sometimes over the course of a patientlife-time, the distal tip electrode to tissue interface increases inimpedance. This means that the new pacemaker would have to pulse at avery high voltage output level which would quickly deplete its batterylife. This is yet another example of why a physician would choose toinsert new leads. Sometimes the old leads are simply extracted. However,this is a very complicated surgical procedure which does involve risksto the patient. Fortunately, there is plenty of room in the venoussystem and in the tricuspid valve to place additional leads through thesame pathway. The physician may also choose to implant the pacemaker onthe other side. For example, if the original pacemaker was in the rightpectoral region, the physician may remove that pacemaker and choose toinstall the new pacemaker in the left pectoral region using a differentpart of the venous system to gain lead access. In either case, theabandoned leads can be very problematic during an MRI procedure. Ingeneral, abandoned leads are capped at their proximal connector pointsso that body fluids will not enter into the lead system, causeinfections and the like. However, it has been shown in the literaturethat the distal electrodes of abandoned leads can still heat up duringMRI procedures. Accordingly, a passive frequency selective circuit ofthe present invention is very useful when placed at or near the proximalelectrical contact after a pacemaker is removed and its leads aredisconnected (abandoned). For example, for an abandoned (left in thebody) lead, an energy dissipating surface 161 c at or near the proximallead end is an ideal place to eliminate excess energy induced by MRI inthe leadwire system. As it will be shown in subsequent embodiments, theenergy dissipating surface 161 can actually be a conductive housing ofthe AIMD 102. Referring back to the article by Dr. Bruce Wilkoff,attention is drawn to slide number 2, which is an example of a leadextraction showing both a distal tip electrode and a distal ring whichhave been heavily overgrown and encapsulated by body tissue. Specialcutting tools were used to free the lead so it could be extracted, sothe tissue shown here is only a small remaining portion of the mass thatwas originally present. Slide 13 is a dramatic illustration of what alarger mass of encapsulated tissue would look like. In this case, theentire tip was completely surrounded, but if one looks carefully to theright, one can see that some of the ring was still exposed. Thesituation is highly variable in that the ring is not always fullyencapsulated. Slide 16 is an example of tissue removal after a pacemakerbipolar lead was extracted. One can see at the end of the lead, thehelix screw that was affixed to myocardial tissue. The surgeon in thisphoto was removing the tissue encapsulation, which completely surroundedthe tip and is still surrounding the ring area. A blowup of this isshown in slide 17. Again, the tissue that is still affixed to the leadhas completely encapsulated the ring, which cannot be seen. Accordingly,there is a need for either a way to prevent the overgrowth of bodytissue onto the ring or to ensure that an energy dissipating surface 161is located far enough away from myocardial tissue to guarantee that itwill remain floating in the blood pool.

FIG. 28 illustrates an energy dissipating ring 161 which is located atsome distance “d” from both a pacemaker tip electrode 138 and a ringelectrode 108 mounted respectively at the distal ends of leadwires 104and 106. The distance “d” should be sufficient so that the energydissipating surface 161 is far enough away from both the distal tip andring electrodes 138, 108 such that there is no heating or temperaturerise associated with the tissues that contact the tip and ringelectrodes. Another advantage of moving the energy dissipating surface161 away from the distal electrodes, particularly for a cardiacpacemaker application, is that there would be less tendency for theenergy dissipating surface 161 to become encapsulated or overgrown withbody tissue. If the energy dissipating surface 161, when it isdisassociated at some distance from the electrodes 138, 108, does becomeovergrown with body tissue, this is not of great concern. Obviously, itwould be superior to have the 161 surface floating in freely flowingblood so that there would be constant cooling. However, for example, ifthe 161 surface did touch off to the right ventricular septum and becameovergrown, the only effect would be a slight heating of tissue in anarea that is far away from where the actual electrical stimulation andsensing is being done by the electrodes. The ideal distance for theenergy dissipating surface does depend on the particular application andranges from approximately 0.1 cm to 10 cm distance from the distalelectrodes.

Referring once again to FIG. 28, the energy dissipating surface 161 isshown as a cylindrical ring. It can be semi-circular, rectangular,octagonal, hexagonal or even involve semi-circles on the lead or anyother metallic or similar structure that is also thermally conductive.Literally any shape or geometry can be used as an energy dissipationsurface 161. It is a desirable feature of the present invention that thesurface area be relatively large so that it can efficiently dissipateheat into the surrounding blood pool and surrounding tissues that aredistanced from the electrodes. In FIG. 28, within the ring 161, thereare electrical connections (not shown) between leadwire 104 and 106 andto the energy dissipating surface 161 that embody the passive frequencyselective circuits previously discussed in connection with FIGS. 2through 11. The purpose of these frequency selective circuits is toremove RF induced energy caused by the RF pulsed field of MRI fromleadwires 104 and 106 and redirect it to the surface 161 where it isdissipated as heat. By having a large surface area, the heat can bedissipated without significant temperature rise such that surroundingtissues would be burned.

In cardiac rhythm management applications, the surface 161 is ideallylocated in areas where there is freely flowing blood, lymph orequivalent body fluids which adds to the cooling. A biomimetic coating154 can be applied to the energy dissipating surface area 161 and/or tothe ring electrode 108 if it is used as an energy dissipating surface161. This special biomimetic coating 154 provides a non-thrombogenic andanti-adhesion benefits. This coating can be comprised of a surfactantpolymer having a flexible polymeric backbone, which is linked to aplurality of hydrophobic side chains and a plurality of hydrophilic sidechains. This coating prevents the adhesion of certain plasma proteinsand platelets on the surface and hence initiation of the clottingcascade or colonization of bacteria. Biomimetic coatings also tend toprevent overgrowth or adhesion of body tissues as illustrated in theWilkoff paper. This polymer compound is described in U.S. Pat. No.6,759,388 and U.S. Pat. No. 7,276,474, the contents of both patentsbeing incorporated by reference herein. Additional benefits ofbiomimetic coatings include the prevention of bacterial colonization andresulting infections. It will be obvious to those skilled in the artthat other types of coatings could be used on the 161 ring to inhibit orprevent overgrowth of body tissue. As used herein, the term biomimeticsincludes all such type coatings.

FIG. 29 is a typical quad polar neurostimulation lead system. It will beappreciated that the following discussion also applies to bipolar, hexpolar, and even sixteen to twenty-four electrode lead systems (thepresent invention is applicable to any number of implanted leads orleadwires or electrodes). In FIG. 29, four leadwires 104 a, 104 b, 106 aand 106 b are shown which are each directed respectively toward anassociated distal electrode 108 a, 108 b, 110 a and 110 b. In this case,the electrical stimulation pulses are applied in various combinationsbetween the various electrodes. Unlike a cardiac pacemaker application,there is no particular ring electrode in this case. However, theinsulation sheath 142 that surrounds the leadwires, which as mentionedcould be of silicone or the like, forms a surrounding surface, whichencapsulates the leadwires.

Parasitic capacitances 22 are formed respectively between each of theleadwires 104 a, 104 b, 106 a and 106 b and the insulating sheath 142.As previously mentioned, these parasitic capacitances are desirable asthey divert high frequency pulsed RF energy from an MRI system to theinsulation sheath 142 thereby redirecting the energy so that heat willbe dissipated over a larger surface area and away from the interfacebetween the distal tip electrodes 108 a, 108 b, 110 a, and 110 b andbody tissue. There is also heat that is directly dissipated off of theleadwires, which is conductively coupled into the insulation sheath 142.Again, it is desirable that this occur at a location that is spaced fromor distant from the therapy delivery electrodes 108 a, 108 b, 110 a, and110 b. This can be greatly improved by providing a passive componentfrequency selective diverter circuit 112 which provided a very lowimpedance at a selected high frequency or frequencies between each ofthe associated leadwires and the energy dissipating surface 161. Theenergy dissipating surface 161 would typically either be a metallic ringor a metallic plate or even a separated metallic surface which has boththe property of conducting the high frequency energy and also having arelatively large surface area for dissipating said energy intosurrounding body tissues. In a preferred embodiment, the energydissipating surface 161 would be placed sufficiently far back from thedistal electrodes 108 a, 108 b, 110 a, and 110 b so that in theassociated heating of surrounding body tissue would not have any effecton the delicate electrode-to-tissue interface. In addition, by having anenergy dissipating surface 161 with a sufficiently large surface area,this will prevent a dangerously large temperature rise as it dissipatesenergy into the surrounding tissues. By controlling the temperature riseto a small amount, damage to tissue or tissue changes are thereforeavoided. The frequency selective reactances 112 are designed to presenta very low impedance at selected high frequencies thereby redirectingundesirable high frequency RF energy (in the MHz range) away from theelectrodes to the insulating sheath and/or energy dissipating surface161. In addition, further protection is offered by the optional seriesfrequency selective components 118 and 120. Typically, these can beseries inductors or they can be parallel inductor capacitor bandstopfilters in accordance with the present invention (see FIGS. 10-11).Accordingly, substantial protection is provided such that during MRIprocedures, the distal electrodes 108 a, 108 b, 110 a, 110.sub.n do notoverheat.

FIG. 30 is taken from FIG. 13 of U.S. 2008/0132987 A1, the contents ofwhich are incorporated herein by reference. Illustrated is a side viewof the human head with a deep brain stimulation electrode shaft assembly160. At the distal end of the electrode shaft 160 are two distalelectrodes 108 and 110 (see FIG. 31) implanted into the patient's brainmatter 162 at a selected implantation site (there can be any number ofelectrodes). One or more leadwires 104, 106 (see FIG. 27A) are routedbetween the skin 164 and the skull 166 down to a pectorally implantedAIND (pulse generator) which is not shown. Referring back to FIG. 30,one can see that a burr opening 168 in the skull 166 has been made sothat the electrode shaft assembly 160 can be inserted.

FIG. 31 is taken generally from section 31-31 in FIG. 30. Shown arebipolar distal electrodes 108 and 110 at or near the end or tip 170 ofthe electrode shaft 160. The skull is shown at 166 and the dura is shownas 172. Housing 174 acts as an energy dissipating surface 161 and can behermetically sealed to protect the passive frequency selective diverterand/or impeder components of the present invention from direct exposureto body fluids.

FIG. 32 is taken from section 32-32 of FIG. 31. Shown are frequencyselective passive component diverter circuit elements 112 which aregenerally taken from FIG. 5 or 6. As previously described, thesediverter circuit elements 112 could be combined with series impederreactance elements 118 and 120 as previously illustrated in FIGS. 7, 10and 11. These have been omitted for clarity, but would generally beplaced in series with the leadwires 104 and 106 and placed betweenfrequency selective circuit elements 112 and the distal electrodes 108,110. Referring back to FIG. 32, circuit elements 112 would divert highfrequency RF energy induced from an MR scanner to the energy dissipatingsurface 161 where it would be dissipated as RF or thermal energy intothe area of the skull 166 and/or dura 172. Frequency selective circuitelement 112 b is also shown connected between the leadwires 104 and 106.This is optional and would be effective for any differential modesignals that are present in the leadwires 104 and 106. In accordancewith FIG. 4 of the present invention, the diverter 112 would redirect ordivert MRI induced RF energy back into leadwires 104 and 106 and awayfrom the distal electrodes 108, 110. This is an example of redirectingRF or thermal energy away from a critical tissue interface point. Theskull is considered to be a relatively non-critical or less susceptibletype of body tissue to thermal injury. This is in comparison with thevery thermally sensitive brain matter into which the distal tipelectrodes 108, 110 are implanted. It has been shown that even atemperature rise as small as a few degrees C. can cause damage tosensitive brain matter.

FIG. 33 is generally taken from area 33-33 of FIG. 31. Shown are the twobipolar electrodes 108 and 110. The frequency selective elements 112 and112 b have been moved relative to the location shown in FIG. 32 toillustrate one wrong way to approach this particular problem.Specifically, an energy dissipating surface 161 is shown mountedgenerally at or near the end portion of the probe shaft 170 in proximityto and/or direct contact with sensitive brain tissue. The frequencyselective reactance components 112 and 112 b are coupled for redirectingthe RF energy from MRI to the energy dissipating surface 161, wherebyheat will be dissipated by the energy dissipating surface 161. In thecase where it was chosen not to use an energy dissipating surface 161,but simply to rely on the line-to-line frequency selective element 112b, heat would still build-up in the entire distal electrode area andthence be conducted into thermally sensitive brain tissue 162.Accordingly, the placement of the circuit elements as shown in FIG. 33illustrates a disastrous way to place the frequency selective elementsof the present invention. Severe overheating of this distal tip wouldoccur with resulting brain damage. Reference is made to a paper given atthe 8.sup.th World Congress of the National Neuromodulation Societywhich was held in conjunction with the 11.sup.th Annual Meeting of theNorth American Neuromodulation Society, Dec. 8-13, 2007, Acapulco,Mexico. This paper illustrates severe tissue damage surrounding a distaltip electrode. This paper was given by Dr. Frank Shellock, Ph. D. andwas entitled, MRI ISSUES FOR NEUROMODULATION DEVICES.

Shellock slide 31 shows X-ray views of the placement of deep brainstimulator electrodes into the skull and brain of a human patient. Thereis also an X-ray view showing the placement of the AIMDs and tunneledleadwires that are associated with the deep brain stimulationelectrodes. Slide number 35 shows an extensive thermally induced lesionshown in white with a red arrow to it. This was representative of twopatients that inadvertently received MRI wherein their deep brainstimulators overheated and caused extensive thermal injury to the brain.Both patients had neurologic deficits and were severely disabled.

In summary, the configuration illustrated in FIGS. 30, 31 and 32,wherein the thermal energy as dissipated into the skull or dura, ishighly desirable as compared to the configuration as illustrated in FIG.33, which could cause thermal damage to sensitive brain tissue.

Referring once again to the Shellock paper, one can see that the deepbrain stimulator involved multiple electrodes. In FIG. 31 one can seethat in this example, there are only two electrodes 108 and 110. This isa way of illustrating that with real time MRI guidance, the physiciancan much more accurately place the electrodes into the exact area of thebrain, which needs to be electrically stimulated (for example, tocontrol Parkinson's tremor, Turret's Syndrome or the like). What istypically done is that precise MR imaging is performed prior toelectrode implantation which is referenced to fiducial marks that'splaced on the skin in several locations outside of the patient's skull.The patient's head is first shaved, then these marks are placed and thenthe MRI is performed. Then when the patient enters the operating room, afixture is literally screwed to the patient's head at these fiducialmarks. This fixture contains a bore into which the various drilling andelectrode implanting tools are located. Because of the need for all ofthis mechanical fixturing, tolerances are involved. This means that bythe time the electrodes are implanted in the brain, they may be not inthe precise locations as desired. Accordingly, extra electrodes areinserted which involves more leads than are really necessary. Thepatient is usually awake during parts of this procedure wherein thephysician will use trial and error to stimulate various electrode pairsuntil the desired result is achieved. In contrast, the present inventionminimizes the need for all these extra electrodes and extra wiring. Thisis because by eliminating the potential for the distal electrodes tooverheat and damage brain tissue, this entire procedure can be doneunder real time MRI imaging. In other words, the physician can bewatching the MRI images in real time as he precisely guides theelectrodes to the exact anatomy of the brain that he wishes tostimulate.

FIG. 34 is a hermetically sealed package consisting of a passive distaltip electrode 138 which is designed to be in intimate contact with bodytissue, such as inside the right atrium of the heart. A hermetic seal isformed at laser weld 176 as shown between the tip electrode 138 and ametallic ring 178. Gold brazes 180 are used to separate the metallicring 178 from the energy dissipating surface 161 by use of anintervening insulator 182. This insulator 182 could typically be ofalumina ceramic, other types of ceramic, glass, sapphire or the like.The energy dissipating surface 161 is typically gold brazed 183 to theother side of the insulator 182 as shown. An inductor 116, such as aninductor chip in accordance with FIG. 10, is shown connected between thedistal tip electrode 138 and a conductive leadwire or pin 184 which isattached by laser welds 176 to the end of the leadwire 104 extending tothe AIMD. As shown, the lead 184 protrudes through a hermetic sealassembly 188 formed by a metallic flange 186 which is typically oftitanium or platinum or the like. The flange 186 is hermeticallyattached to the lead 184 by gold brazes 180, and is typically laserwelded as shown at 177 to a proximal end of the energy dissipatingsurface 161.

FIG. 35 is a cut-away view taken generally from the housing of FIG. 34.It is important that the electrical insulating material 182 either be ofvery low thermal conductivity or have a relatively long length “L” asshown. The reason for this is that the thermal energy that is developedin the energy dissipating surface 161 must not be allowed to reach thedistal tip electrode 138 as shown in FIG. 34 where heat could causedamage to the adjacent tissue.

The energy dissipating surface 161 is typically of biocompatible metals,such as titanium, platinum or the like. It is important that the energydissipating surface be both electrically conductive and thermallyconductive so that it can transfer RF and thermal energy into body fluidor tissue. The energy dissipating surface 161 can be roughened or evencorrugated or bellowed as shown in FIG. 36 to increase its surface areaand therefore its energy dissipating properties into Surrounding bodyfluids or body tissue.

In accordance with FIG. 5, capacitive elements 114 are shown in FIG. 34are designed to act as a low impedance at higher frequencies. Electricalconnections 190 (FIG. 34) couple the capacitors 114 from the leadwire184 to the energy dissipating surface 161. This forms a broadband lowpass filter wherein the inductor 116 acts in cooperation with thecapacitive elements 114. The presence of the inductor element 116 is notrequired; however, it does enhance the performance of the capacitorelements 114. Capacitor elements 114 are typical off-the-shelfcommercial monolithic ceramic capacitors (MLCCs). These are betterillustrated in FIG. 38.

There is an advantage in the present invention in using a capacitor forthe selective frequency element 112 as shown in FIG. 5. The capacitortends to act as a broadband filter which will attenuate a range of MRIfrequencies. For example, placement of an effective capacitor 114 couldattenuate 64 megahertz, 128 megahertz and higher MRI frequencies.However, if one were to use an L-C series trap filter as shown in FIG. 6for the variable frequency element 112, then this would only beeffective at one MRI frequency, for example, 64 megahertz only. Ofcourse, as already been disclosed herein, one could use multiple L-Ctrap filters. However, in a preferred embodiment the use of a capacitoras illustrated in FIG. 5 is desirable because with a single component,one can attenuate a broad range of MRI frequencies.

The schematic diagram for the circuitry of FIG. 34 is shown in FIG. 37.Capacitors 114 are actually in parallel and act as a single capacitivediverter element to the EDS surface. The reason for multiple capacitorsis to obtain a high enough total capacitance value so that thecapacitive reactance is very low at the frequency of interest (forexample, 64 MHz for a 1.5 T MR system).

An alternative capacitor 114 for use in the circuit of FIG. 37 is knownas a unipolar feedthrough capacitor is shown in FIG. 39. It has insidediameter and outside diameter termination surfaces 192 for electricalcontact. Feedthrough capacitors can be unipolar or multipolar. These arecompletely described in the prior art; for example, refer to U.S. Pat.No. 7,363,090, particularly FIGS. 3, 5, 29 through 31, and 39. See alsoU.S. Pat. Nos. 4,424,551; 5,333,095; and 6,765,779.

FIG. 40 is similar to FIG. 34 (using common reference symbols) exceptthat the inductor element 116 is wire wound around a non-ferromagneticmandrel 194 (formed from a material such as a ceramic or plastic). Thistype of wound inductor 116 has much higher current handling capabilityas compared to the inductor chip of FIG. 34. The inductor chip of FIG.34 can be fabricated from a variety of shapes including Wheeler'sspirals and the like. Refer to U.S. 2007/0112398 A1, FIG. 83 and FIGS.70 and 71 of U.S. 2009/0243756. These inductors can be manufactured by anumber of printing techniques including lithographic or copper cloutingand etching. However, this results in relatively thin and highresistivity inductor traces.

It is important that the inductor element 116 of the present inventionbe able to handle substantially high currents when it is in series withthe lead 184. The reason for this has to do with either ICD applicationsfor shock electrodes or automatic external defibrillation (AED) events.AEDs have become very popular in government, buildings, hospitals,hotels, and many other public places. When the external defibrillatorpaddles are placed over the chest of a cardiac pacemaker patient, thehigh voltage that propagates through body tissue can induce powerfulcurrents in implanted leads. Accordingly, the inductor 116 of thepresent invention has to be designed to handle fairly high current (ashigh as the 4 to 8 amp range in short bursts). The wire wound inductor116 of FIG. 40 has wire of a larger cross-sectional area and istherefore a higher current handling inductor and is therefore apreferred embodiment.

FIG. 41 illustrates an entirely different approach for the diverting ofRF energy away from the electrode 138 to the energy dissipation surface161. Shown are electrical connections 196 a, 196 b between a firstinductor 116 a and the distal tip electrode assembly 138. The other endof the first inductor 116 a is connected to a second inductor 116 bwhich is in turn electrically connected at 116 c to the leadwire 184,104. The capacitor 114 is connected between the junction of the twoinductors 116 a and 116 b at electrical connection 196 d. The other endof the capacitor is electrically connected at 196 e to the energydissipating surface 161. An insulating sleeve (not shown) can be used toensure that the capacitor termination and electrical connection 196 ddoes not inadvertently make contact (short out) with the energydissipating surface 161. As shown, this connection is made adjacent tothe insulator 182 so there is no chance for such shorting out.

The electrical schematic for FIG. 41 is shown in FIG. 42. In accordancewith FIG. 7, this forms what is known in the art as a low pass filter(in this example, a T filter), which tends to enhance the filteringperformance by directing more of the RF energy to the energy dissipatingsurface 161. As previously mentioned, a single or multi-element low passfilter would attenuate a broad range of MRI frequencies and would be anadvantage in the present invention for that reason.

The various types of low pass filters are more thoroughly shown in FIGS.43 and 44 which compares the filtering efficiency measured asattenuation in dB with increasing numbers of filter elements. The lowpass filters illustrated in FIG. 43 perform two very importantfunctions. First, they are very effective EMI filters in order toprotect AIND electronics from the powerful electromagnetic fields duringMRI scans and the like. Secondly, they all have capacitor diverterelements that are associated with their inductor impeder elements.Accordingly, the capacitors act as energy diverters in the presentinvention thereby redirecting induced RF energy on the leads to theenergy dissipating surfaces (161). Shown are single element low passfilters consisting of either the capacitor 114 or an inductor 116, an Lfilter which consists of an inductor 116 and a capacitor 114, a Tfilter, a Pi filter (FIG. 44), an LL filter (FIG. 44) or an “n” elementfilter (FIG. 43). FIG. 43 shows the general response curves of thesetypes of filters as attenuation versus frequency. Selected schematicsfor these various filters, which are correlated to the curves in FIG.43, are shown in FIG. 44. As one increases the number of filterelements, the ability to attenuate or block high frequency signals fromreaching sensitive AIMD electronics is improved. Referring once again toFIG. 43, for example, one can see that for a particular value of asingle element capacitive filter, the attenuation for a 1.5 Tesla MRIsystem operating at 64 MHz is only about 12 dB. This means that acertain amount of the RF energy would still reach the distal tipelectrode. Now compare this to the T filter of. FIG. 43, where one cansee that there is in excess of 45 dB of attenuation. In this case, aninsignificant amount of RF energy from the RF pulsed frequency of theMRI, would reach the distal electrode. Accordingly, one preferredembodiment of the present invention is that a capacitor combined withone or more inductors would be an optimal configuration. As the numberof elements increases, the filtering efficiency improves. When thefiltering efficiency improves, this means that less and less RF energywill reach the distal tip.

FIG. 45 illustrates a schematic diagram of a series inductor116-capacitor 114 filter which is commonly known in the industry as anL-C trap filter. The L-C trap filter was previously described inconnection with FIG. 6. Referring once again to FIG. 45, there is aparticular frequency for a trap filter when the capacitive reactancebecomes equal and opposite to the inductive reactance. At this singlefrequency, the capacitive reactance and the inductive reactance canceleach other out to zero. At this point, all one has left is the residualresistance 198. If one selects high quality factor (Q) components,meaning that they are very low in resistance, then the trap filter ofFIG. 46 ideally tends to look like a short circuit at its resonantfrequency f, between points A and B which may comprises connectionsrespectively to a pair of leadwires 104 and 106. FIG. 46 gives theresonant frequency equation where f_(r), in this case, was measured inhertz. FIG. 9 shows the effect of a short circuit 122 between leadwires104 and 106. Referring once again to FIG. 45, it is important that theamount of resistance 138 be controlled. This is better understood byreferring to FIG. 47.

FIG. 47 illustrates the impedance Z in ohms versus frequency of theseries resonant L-C trap filter of FIG. 45. As one can see, theimpedance is quite high until one reaches the frequency of resonancef_(r). At this point, the impedance of the series L-C trap goes very low(nearly zero ohms). For frequencies above or below resonance f_(r),depending on the selection of component values and their quality factor(Q), the impedance can be as high as 100 to 1000 or even 10,000 ohms orgreater. At resonance, the impedance tries to go to zero and is limitedonly be the amount of resistance 138 (FIG. 45) that is generallycomposed of resistance from the inductor 116 and also the equivalentseries resistance from the electrode plates of the capacitor 114. Theresistance 138 could also be a discrete resistor that is added in serieswith the capacitor 114 and the inductor 116. In a preferred embodiment,this would be a chip resistor. There is a trade off in proper selectionof the components that controls what is known as the 3 dB bandwidth. Ifthe resistance is extremely small, then the 3 dB bandwidth will benarrower. However, this makes the trap filter more difficult tomanufacture. Accordingly, the 3 dB bandwidth and the resistive element Rare preferably selected so that it is convenient to manufacture thefilter and tune it to, for example, 64 MHz while at the same timeproviding a very low impedance R at the resonant frequency. For an idealL-C series resonant trap filter, wherein ideal would mean that theresistance R would be zero, then the impedance at resonance would bezero ohms. However, in this case, the 3 dB bandwidth would be so narrowthat it would be nearly impossible to manufacture. Accordingly, someamount of resistance R is in fact desirable. In a preferred embodiment,the resistance R in an L-C trap filter is equal to the characteristicresistance of an implanted lead so that maximum energy transfer willoccur to the energy dissipating surface. The resistance in the L-Cseries resonant trap filter can be controlled by the amount ofresistance in the inductor itself and/or the equivalent seriesresistance of the capacitor. For example, one can control or evenincrease the resistance of an inductor by adding more turns of wire ormaking the wire or circuit traces smaller in cross-section. One couldalso use higher resistivity materials in the construction of theinductor. One could also add a discrete resistor in series with theinductor and capacitor of the L-C trap. In a preferred embodiment, thiscould be a discrete chip resistor. In summary, controlling theresistance of the L-C trap filter is a novel feature in its applicationherein as a frequency selected diverter to a surface 161.

As previously mentioned, there is a disadvantage to use the L-C trapfilter as shown in FIG. 6. That is, it is really only effective forattenuating the one MRI frequency (for example, 64 megahertz for a 1.5megahertz scanner).

Accordingly, when the AIMD manufacturer would apply for their FDAconditional labeling, they could only claim compliance with 1.5 TeslaMRI scanners. However, the L-C trap filter of FIG. 6 also offers a veryimportant advantage in that it offers a very high degree of attenuationat this one selected frequency and is also highly volumetricallyefficient. Accordingly, there is a tradeoff here. When one uses abroadband low pass filter, a broad range of frequencies is attenuated atthe cost of increased size and complexity (an additional number ofcomponents). An L-C trap filter such as shown in FIG. 6 is more of a“rifle-shot” approach wherein only one selected frequency is attenuated.In physics, this is more efficient and tends to make the componentssmaller. By controlling the value of the resistance 138 in FIG. 45,energy transfer is maximized from the implanted leads to the 161 surfaceor housing of the AIMD. In accordance with Thevenin's Maximum PowerTransfer Theorem, assuming a resistive system, maximum energy transferto a load occurs when the characteristic source impedance (the leadsystem impedance) is equal to the load resistance. For example, if theimplanted lead had an implanted characteristic resistance of 2 ohms, itwould be desirable to have the resistance of the L-C trap filterillustrated in FIG. 45 also be 2 ohms. A potential disadvantage of ahigh Q (low resistance) L-C trap filter is that at resonance, itsinductive and capacitive reactive components cancel each other out. Inother words, at resonance the L-C trap becomes purely resistive. Inaccordance with the present invention, it is a relatively simple matterthough to add resistance in series with the L-C trap filter. This can bedone through using a discrete resistor such as a chip resistor or bydeliberately building addition resistance into the design of theinductor and/or the capacitor's equivalent series resistance.

FIG. 48 illustrates yet another method of decoupling RF signals fromleadwire 104. Referring back to FIGS. 34 through 43, all of theaforementioned decoupling techniques involve broad band low passfiltering. The advantage with these is that they would be applicable toa wide range of MRI machines including 0.5, 1.5, 3.0, 5.4 Tesla and soon. In other words, these broad band EMI filters would attenuate a broadrange of RF frequencies. In FIG. 48, one can see that there are twodiscrete L-C trap filters. The first trap filter consists of inductor116 a and capacitor 114 a acting in series, and the second trap filterconsists of inductor 116 b and capacitor 114 b operating in series. Thisis best understood by referring to the schematic of FIG. 49 which showsthe series connection of 116 a, 114 b from the lead 184 to the energydissipating surface 161. Inductor 116 b and capacitor 114 b are alsoconnected in series from the lead 184 to the energy dissipating surface161.

In FIG. 48, one can see that an electrical connection 196 a is madebetween the distal tip electrode 138 and inductor chip 116. Inductorchip 116 a is then electrically connected via electrical connectionmaterial 196 b to monolithic chip capacitor (MLCC) capacitor 114 a. Theother end of the chip capacitor 114 a is electrically connected at 196 dto the energy dissipating surface 161. Inductor 116 b is also connectedto the distal tip electrode 138 by material 116 e. The other end ofinductor 116 a is connected in series at 116 c with capacitor 114 b. Theother end of capacitor 114 b is electrically connected at 116 f to theenergy dissipating surface 161. In this way, the two trap filters areconnected in parallel between the lead 184 and the energy dissipatingsurface 161 as shown in the schematic diagram of FIG. 49.

FIG. 50 illustrates a typical chip inductor 116 a, 116 b which can beused in FIG. 50.

FIG. 51 is a typical prior art MLCC chip capacitor 114 a, 114 b whichcan also be used in conjunction with the package shown in FIG. 48.

FIG. 52 is a graph of impedance versus frequency showing the impedancein ohms for the two L-C trap filter elements that were previouslydescribed in FIGS. 48 and 49. By carefully selecting the componentvalues 114 a and 116 a and also 114 b and 116 b, one can select thefrequencies at which the two (or more) L-C trap filters willself-resonate. In the present example, the first trap filter includingcomponents 114 a and 116 b has been selected to resonate at 64 MHz, andthe second trap filter including element 114 b and 116 b has beenselected to resonate at 128 MHz.

Referring once again to FIG. 52, one can see that we now effectivelyhave dual trap filters which tend to act as very low impedance betweenthe leadwire 184, 104 and the energy dissipating surface EDS at twodifferent frequencies. In this case, by example, the first trap filterresonates at 64 MHz, which is the RF pulsed frequency of a 1.5 Tesla MRIsystem. The second trap filter, which has resonant frequency 128 MHz, isdesigned to divert RF energy to the EDS surface from a 3 Tesla MRIfrequency. It will be appreciated that a multiplicity of trap filterscan be used depending on how many different types of MRI systems thatone wants to claim compatibility with for an implanted lead andelectrode. The method of selecting the resonant frequency was alreadydescribed in FIG. 46 and is applicable to FIG. 52. Referring once againto FIG. 52, one will note that except at the resonant frequency f.sub.r1and f.sub.r2, the impedance of the trap filter is very high. This isvery important so that low frequencies are not attenuated. Accordingly,using a cardiac pacemaker application as an example, pacing pulses wouldbe free to pass and also low frequency biologic signals, such as thosethat are produced by the heart. It is very important that pacemakersensing and pacemaker pacing can occur while at the same time, highfrequency energy, for example, that from the RF pulsed frequency of anMR system can be diverted to an appropriate energy dissipating surface161. The parallel trap filters, as described in FIGS. 48, 49 and 52,have to be carefully designed so that they will not resonate or interactwith each other. This is best accomplished if one were to place abandstop filter between them which would tend to electrically isolatethem. This is not shown, but would be understood by those skilled in theart.

FIG. 53 illustrates a typical active implantable medical device bipolarleadwire system. On the left is shown a distal tip electrode 138 and adistal ring electrode 108. The energy dissipating surface 161 of thepresent invention is shown along with coaxial leadwires 104 and 106which would be connected to the AIMD. These could be endocardial orepicardial in accordance with the prior art.

FIG. 54 is a blown up sectional view generally taken from section 54-54from FIG. 53. In FIG. 54, one can see that there is an energydissipating surface 161 which is enclosed at both ends by two hermeticseal flanges or flange assemblies each consisting of a flange 188, aninsulator 200 and gold brazes 180, 180′. The flange 188 is designed tobe laser welded 176 into the metallic energy dissipating surface 161 asshown. A bipolar feedthrough capacitor 114 c is shown in cross-sectionin FIG. 54 where the two leadwires 104 and 106 pass through its Thefeedthrough capacitor 114 c is a very efficient broadband filter whichwould tend to decouple or divert high frequency signals such as 64 MHz(1.5 Tesla) and 128 MHz (3 Tesla) from the leadwires 104, 106 to theenergy dissipating surface 161 in accordance with the present invention.Each leadwire 104 and 106 may additionally include the frequencyselective impeding reactances 118 and 120 (as previously shown anddescribed in FIGS. 7, 10 and 11).

The bipolar feedthrough capacitor 114 c is illustrated in isometric viewin FIG. 55. Shown is an outside diameter termination surface 192 b whichis electrically and thermally connected to the inside diameter of theenergy dissipating surface 161 of FIG. 54, as by electrical connection196 g. Also shown, are termination surfaces 192 a 1 and 192 a 2 locatedon the inside diameter of two feedthrough capacitor ID holes forelectrical connection at 196 h and 196 i (FIGS. 54, 55) betweenleadwires 104 and 106, respectively to the feedthrough capacitortermination surfaces 192 a 1 and 192 a 2, respectively. The use of afeedthrough capacitor in this case makes for a truly broadbandperformance. As MR systems continue to evolve in their static magneticfield strength, the RF pulse frequencies go higher and higher. Forexample, for a 10 Tesla scanner, the RF pulse frequency is 426.5megahertz. Prior art MLCC chip capacitors have internal inductance andtend to self-resonate at frequencies around 400 megahertz or above.Accordingly, the use of a feedthrough capacitor accommodates much higherfrequency MRI systems.

Referring once again to FIG. 28 and FIG. 31, one can understand why theenergy dissipating surface 161 of FIG. 53 has been moved back a suitabledistance “d” from the distal tip electrode 138 and the distal ringelectrode 108. This is because of the tendency for distal tip 138 andring electrodes 108 to become completely embedded or encapsulated withbody tissue. In other words, one cannot guarantee that the distal ringelectrode 108 will always be freely floating in the blood pool, forexample, of the right ventricle or the right atrium. Referring onceagain to FIG. 27, one can see shaded areas where tissue encapsulationtends to be the greatest. An ideal location for the energy dissipatingsurface 161, as described in FIG. 53, is shown as 161 in FIG. 27. Thisguarantees that the energy dissipating surface 161 is placed generallyinto the area of the right ventricle that is free of trabecula tissueand where there is always freely flowing blood. Of course, this isparticularly important for cardiac rhythm management applicationswherein pacemakers and implantable defibrillators are commonly used. Forimplantable neurostimulators, generally, these are not placed in areaswhere there is freely flowing blood. However, it is still important inthese cases that the energy dissipating surface 161 be a sufficientlylarge enough distance from the associated electrode(s) so that if thereis adjacent tissue heating, it does not affect the delicate interfacebetween the electrodes and surrounding body tissue. This would beparticularly important, for example, in a deep brain stimulator. Asshown in FIG. 31, for example, an ideal location for the energydissipating surface 161 would be either at the skull or subdural(slightly below the skull). In this case, the deep brain stimulationelectrode would protrude down into the brain tissue below the energydissipating surface 161. In this case, the RF energy and/or heat wouldbe dissipated over a relatively large surface area well away from thevery heat sensitive and delicate brain tissues 162. For a spinal cordstimulator, there is generally freely flowing spinal fluid which can actas a cooling agent as well. In this case, it is desirable to have thesurface 161, again, spaced at some distance from the therapy deliveryelectrode such that cooling effectively takes place within the cerebralspinal fluid. See U.S. Publication Nos, US 2008/0132987 A1 and US2007/0112398 A1, which are incorporated by reference herein. In somecases, the separation distance can be quite small, for example on theopposite surface of a paddle electrode as shown herein in FIGS. 72, 73and 74.

FIG. 56 is a schematic diagram of the energy dissipating surface 161assembly previously described in FIGS. 53 and 54. In FIG. 56, one cansee that the passive frequency selective diverter elements 114 a and 114b could be replaced by any of the circuits previously described in FIGS.4 through 11 as element 20.

FIG. 57 illustrates a bipolar lead of the present invention with distaltip and ring electrodes 108, 110 (not shown) at a suitable distance dfrom an energy dissipation surface 161 such that energy dissipation inthe surface 161 would not cause a temperature rise at the distalelectrodes. Shown is a capacitor 114 connected between the leadwires 104and 106. Also shown are a pair of bandstop filters 117 a and 117 b aspreviously illustrated in FIG. 11. Referring once again to FIG. 57, onecan see that the capacitor element 114 acts as a high frequency energydiverter. This works in cooperation with the two bandstop filterelements 117 a and 117 b which act as energy impeders at a selected MRIfrequency. Accordingly, high frequency energy that is induced on theleadwires 104 and 106 is converted to RF circulation currents I₁ and I₂.I₁ and I₂ are shown in opposite directions to illustrate, for example,for a 1.5 Tesla MRI system, that these oscillate back at 64 milliontimes per second. This creates a great deal of current in the associatedleadwires to the right (as viewed in FIG. 47) of the diverting element114. This causes heat to be dissipated in the leadwires 104 and 106 intothe energy dissipating surface 161 such as the overall insulation sheathor shield of the probe, catheter or implanted device as shown.

FIG. 58 is very similar to FIG. 57 except that diverting element 112 hadbeen replaced by a pair of capacitor elements 114 a and 114 b whichconnect from leadwires 104 and 106 respectively to an electromagneticshield or an energy dissipating surface 161. It is a desirable propertyof the present invention that the surface 161 be highly thermallyconductive, have relatively high surface area for efficient transfer ofRF or heat energy into surrounding fluids and body tissue and also beelectrically conductive at RF frequencies. Referring once again to FIG.58, the diverter elements 114 work best when they are on the body fluidside (towards the distal electrode) related to the bandstop filters 117a and 117 b. When the impeders or bandstop filters are placed betweenthe distal electrodes and the impeder capacitors 114, the distalelectrodes will heat up significantly. This is because the energy is nowtrapped in the lead system and reflects back and forth along theimplanted lead which causes distal tip overheating. The bandstop filters117 a and 117 b, if placed incorrectly between the impeder capacitorsand the distal electrodes, represent a very high impedance. This makesit very hard or even impossible for the RF energy entrapped in the leadsystem to reach the surface 161 or housing of the AIMD where it can bedissipated over a large surface area. Instead, the energy bounces backoff and reflects back down to the distal electrodes where itconcentrates as an RF current which causes significant overheating.

The bandstop filters 117 a and 117 b of FIG. 58 look like a very highimpedance (ideally an infinite impedance) at the resonant frequency. Inpractice, the impedance of the bandstop filters at resonance will bearound 2000 ohms. This has the effect of disconnecting or impeding RFcurrent to the distal electrodes at these high frequencies from theleadwires 104 and 106. These work in conjunction with the low passfilter elements 114 a and 114 b which act as a way to divert the highfrequency energy to the energy dissipating surface 161 which in apreferred embodiment is the AIMD housing. As previously mentioned, thelow pass filter elements 114 a and 114 b can consist of any of the lowpass filters as previously described in FIGS. 43 and 44 or the L-C trapfilter as previously described in FIGS. 45, 46, 47, 48, 49 and 52. Ahigh frequency model of FIG. 50 is illustrated in FIG. 9 wherein theleadwires are effectively shorted together to an energy dissipatingsurface 161 and the distal electrodes 108 and 110 have been effectivelycut or disconnected (in this case, by the bandstop filter elements) fromthe electrodes. For a more complete description of bandstop filterelements and their design and operation, refer to U.S. Pat. No.7,363,090.

FIG. 59 illustrates an exemplary bandstop filter 117 a or 117 bconsisting of a parallel inductor 116 and capacitor 114 (as previouslyshown and described herein) with nonlinear circuit elements such asdiodes 202 a and 202 h placed in parallel therewith. These diodes 202 a,202 b are oriented in what is known in the prior art as a back-to-backconfiguration. The diode elements 202 a, 202 b, as illustrated in FIG.49, can be placed in parallel with each other, and with any of thefrequency selective circuit elements as previously described in FIGS. 4through 11. For example, referring to FIG. 5, the diode elements 202 aand 202 b could be placed in parallel with the capacitive element 114.Referring to FIG. 10, two diode elements 202 a, 202 b could also beplaced in parallel with each of the inductor elements 116 a and 116 b.Back-to-back diodes are one form of a transient voltage suppressor.Transient voltage suppressors (TVS) are well known in the prior art forproviding over voltage circuit protection. They are sold under varioustrade names including the name Transorb. The diodes 202 a, 202 b canalso be pin diodes. As previously discussed, automatic externaldefibrillators (AEDs) have become very popular in the patientenvironment. Accordingly, implanted leads must be able to withstand veryhigh pulsed currents. These pulse currents can range anywhere from 1 to8 amps. It is also a feature of the present invention that the passivefrequency selective components be very small in size. In order for aninductor element L to be able to handle 1 to 8 amps, it would have to beexceedingly large. However, by using physically small diode elements 202a and 202 b, one can have the circuits switched to a different state.That is, when a high voltage, such as that from an AED appears, thediodes would forward bias thereby temporarily shorting out the bandstopfilter 117 a or 117 b consisting of the parallel combination of inductorL and capacitor C (FIG. 59). Thereby the correspondingly high AEDinduced currents would be diverted away from the relatively sensitive(small) passive elements L and C in such a way that they not be harmed.

FIG. 60 is nearly identical to FIG. 58 except that transient voltagesuppressors 204 a and 204 b have been added respectively in parallelwith the bandstop filter elements 117 a and 117 b. Transient voltagesuppressors are nonlinear circuit elements which operate in much thesame fashion as previously described for the back-to-back diodes 202 aand 202 b of FIG. 51. This family includes diodes, zener diodes,Transorbs™, Transguard®, metal oxide varistors, Z_(n)0 varisters, andother similar nonlinear circuit elements. The purpose of the transientvoltage suppressors 204 a and 204 b in FIG. 60 is to bypass any highvoltage induced currents such that these currents not flow through therelatively sensitive bandstop passive component inductor and capacitorelements.

FIG. 61 illustrates a general diverter and/or impeder filter element 206which can be representative of any of the filters previously described.The filter element 206 of FIG. 61 is shown disposed between anelectrical connection to an energy dissipating surface 161 which can bean AIMD housing as illustrated. The filter is shown connected to aproximal end of a leadwire 104 or the like with dashed lines, andconnected to a distal end electrode 108 shown coupled to the leadwire104 or the like with dashed lines. The reason for the dashed lines is anindication that the filter 206 can be placed anywhere between the distalend and the proximal end of the leadwire 104 or even inside the AIMDhousing. The filter 206 and energy dissipating surface 161 could belocated near the distal end, at the distal end, at a distal ringelectrode 108 or near a distal ring electrode 108 such that it wouldfloat in the blood pool. The filter 206 can also be placed at or nearthe proximal end, or at any point between the distal and proximal ends.

In particular, the filter and associated energy dissipating surface 161could be located all the way at the proximal end of an abandoned lead.Leads are often abandoned and left in patients for various reasons.Sometimes the lead becomes slightly dislodged, for example, from cardiactissue such that the pacing threshold increases or is lost. Sometimeslead insulation becomes abraded and/or the leadwire itself is broken.Removing leads once they've been in the body for a long time can be verydifficult as portions of the lead tend to become overgrown by bodytissue. One is again referred to the article entitled, ICD EXTRACTIONINFECTED/REDUNDANT LEADS EVERYDAY CLINICAL PRACTICE by Dr. BruceWilkoff. When one looks at the photographs of the extracted leads, onecan see that they are very often substantially overgrown with tissue.Therefore, it is common practice to simply abandon leads.

In the prior art, the abandoned lead is simply capped such that bodyfluid will not enter it. This cap is nothing more than an insulativecap. However, it is also well known in the literature that abandonedleads can be quite dangerous in an MR scanning situation. High energyelectromagnetic fields from the RF pulsed energy of a scannerintensifies at the ends of implanted leads. Because they are abandonedor capped at one end, this creates a reflection situation whereby all ofthe intense energy has no way to escape the lead except at the distalelectrode end. This is the worst case situation because the distalelectrode makes intimate contact with body tissue. For example, if thetissue was myocardial tissue, one runs a severe risk of creating burningor lesions in the heart. In the case of a deep brain stimulator, oneruns the risk of causing deep lesions within the brain. In an abandonedlead, therefore, it is much more desirable that energy be dissipated ator near the proximal end as opposed to the distal end where there aresensitive body tissues involved. In general, active implantable medicaldevices are implanted in muscle or in fat tissues, for example, in thepectoral areas which are not so heat sensitive, but more importantly,are not implanted in an organ, whose function could be compromised.Accordingly, it is a feature of the present invention that any of thefilter networks, as previously described herein, including those asshown in FIGS. 4 through 11, could be incorporated in a cap structure tobe attached to the proximal end of the leadwire wherein such said capstructure includes an energy dissipating surface 161. For a furtherdescription of the problem and the need to provide a cap for abandonedleads, one is referred to U.S. Pat. No. 6,985,775.

FIG. 62 shows an energy dissipating surface 161 in a relatively fixedlocation along the length of a leadwire 104. In accordance with thepresent invention, the energy dissipating surface 161 is placed asuitable distance d from a distal electrode 108 such that energydissipation in the area of the 161 surface will not cause tissueoverheating at or near the distal electrode 108. Also shown is afrequency impeding element 118 which can be moved to various locationsalong the length of the leadwire 104 as indicated by the multipledashed-line boxes 118. For example, impeding element 118 could be placednear the energy dissipating surface 161, or it could be moved toward thedistal electrode 108 at any one of several successive locations. Theimpeding element 118 such as a bandstop filter 117 or a series inductorwill still work in conjunction with the diverting element 112 at any ofthese various locations. In fact, this can be an advantage in thepresent invention in order to make the distal tip electrode 108 and itsassociated leadwire 104 within the distance “d” smaller in diameter. Ingeneral, most leads for cardiovascular applications are restricted tothe six French (0.079 inches in diameter) region. This can beproblematic for a biventricular implant where the endocardial electrodemust be threaded through the venous system and then into the coronarysinus and through the great cardiac vein to one of many branch vesselswhich are outside of the left ventricle. These branch vessels tend to bevery small in diameter and very difficult to navigate, particularly fora large lead (size four French or smaller would be ideal). There is alsoa similar need for certain spinal cord and deep brain stimulators whichmust embody electrodes that are very small in diameter. Referring backto FIG. 62, one can see that by having a relatively large valvecapacitive diverter element 112 associated with a energy dissipatingsurface 161 that is located at a distance d from the distal electrode,one can then downsize the diameter of the wiring along the length ofdistance d. By putting the frequency impeding element such as any one ofthe elements 118 a, 118 b and/or 118 c, one can make this singlecomponent smaller than multiple components. Accordingly, frequencyimpeding elements do not have to be in direct physical proximity todiverting frequency selective elements 112. As taught in FIGS. 4, 5, 6,42 and 43, the diverting element 112 can consist not only in a capacitoror an L-C resonant trap filter, but also could include a variety of lowpass filters. Referring to FIG. 43, for example, one could see that an Lsection low pass filter is identical to the filter described in FIG. 62,wherein element 118 represents the inductor element and element 112represents the capacitor element. Referring once again to FIG. 62, onecan incorporate a T-type filter which embodies two inductor elements. Inthis embodiment, the left hand inductor element 118 would be to the leftof the frequency diverting element 112 and a second inductor (not shown)would be located to the right of the diverter element 112. This righthand inductor could be located in close physical proximity to thediverter element 112, or it could also be moved away as was describedfor the left hand inductor element at various locations as shown in FIG.52.

Referring back to FIG. 62, it should be noted that the variableimpedance element 112 can be monolithic ceramic (MLCC) capacitors,ceramic feedthrough capacitors, or other types of capacitive circuitcomponents. In addition, the frequency selective element 112 can be aparasitic or distributive capacitor wherein the capacitance is formedthrough relatively high-dielectric materials between leadwires orelectrodes in an energy dissipating surface 161.

FIG. 63 illustrates a type of probe or catheter 102 which is typicallyused to both map and ablate the inside of cardiac chambers to eliminateor control certain types of arrhythmias. For example, in a patient withuncontrollable atrial fibrillation, this type of probe or catheter 102would be inserted so that electrical mapping, between bipolar distalelectrodes 108 and 208 or between electrodes 110 a and 110 b, could beperformed to isolate and locate those areas from which the sporadicelectrical activity is occurring. For example, this might be around apulmonary vein. Reference is made to U.S. Pat. No. 7,155,271 for a morecomplete description of this type of need and procedure. After the areasthat need to be ablated are located, the surgeon can apply RF ablationenergy at a distal ablation electrode 208. This has the effect ofburning the inside of cardiac tissue creating a scar which will isolatethis area of erratic electrical activity. The goal here is to complete ascar structure such that the atrial fibrillation is terminated.Unfortunately, in the prior art, this procedure is done using real-timeX-ray, fluoroscopy, landmarks based on CT scans, or other types ofguidance, which does not adequately visualize soft tissue. Accordingly,the surgeon is working pretty much blind as the scars forming cannot beseen in real time. As explained in U.S. Pat. No. 7,155,271, it would bea great advantage if such procedures could be performed during real timeMRI guidance. The problem is the MRI RF energy induced into the ablationcatheter could cause overheating and sporadic formation of scar tissueat the wrong time and/or in the wrong location. In FIG. 63, one can seethat there is a novel energy dissipating surface 161 of the presentinvention. This surface 161 is located at a distance back from thedistal tip such that the energy dissipating surface 161 will redirectenergy away from both the electrical sensing electrodes 108, 110 and theRF ablation electrode 208 where they cannot overheat, at inappropriatetimes. Frequency selective passive components (not shown), in accordancewith the present invention, are connected in series with the leadwires,or from the inside of the energy dissipating surface 161 to the variousleadwires 104, 106 and 210. These are the circuits that have generallybeen described in FIGS. 4 through 11 herein. For simplicity, they havenot been shown in FIG. 63, but should be obvious to one skilled in theart from the previous drawings. In other words, the RF ablationelectrode tip 208 will only overheat when the surgeon decides toactivate the RF circuitry to deliberately form the scar tissue.

The energy dissipating surface 161 may include some materials or antennastructures that are readily visualized during active MRI guidance. Thismay be important so that a physician can ensure that if the probe orcatheter is manipulated that the surface 161 not rest against the insideof, for example, the atrial septum. This is the area that is dissipatingRF energy and heat during the active MRI. If the surface area of thissurface 161 is sufficiently large so that very little temperature risewould occur, it would not matter if the surface 161 touched off against,for example, the inside wall of the cardiac septal wall. However, if thesurface 161 was relatively small, then substantial temperature risecould occur if it was not kept within the freely flowing blood stream.In this case, it would be important that the physician be able tovisualize the surface 161 and the MRI images so that it not be allowedto rest inappropriately against sensitive tissues on the inside of theatrium and cause inadvertent scar tissue or ablation to occur. Referringonce again to FIG. 63, one can see that the ablation electrode 208 isconnected to an RF ablation leadwire 210 which comes from RF ablationequipment (not shown) which is external to the patient. The sensing ringelectrodes 108 and 110 are coupled to leadwires 104 and 106 which runthrough the center of the probe or catheter and also are connected toexternal equipment which is used to monitor electrical cardiac activity.These would typically be connected to an ECG or ERG recorder.

FIG. 64 shows a probe or catheter similar to that illustrated in FIG. 63except that the energy dissipating surface 161 has been convoluted sothat its surface area has been increased. Such increasing of the surfacearea, which is in contact with fluids, such as body fluids, willincrease the amount of MRI induced RF energy that is dissipated.

FIG. 65 is very similar to FIG. 64 except that instead of convolutions,fins 212 have been added. These fins 212 also increase the surface areaand increase the amount of energy or heat which is dissipated intosurrounding fluids and tissues.

FIG. 66 is similar to FIGS. 64 and 65 except that the energy dissipatingsurface 161 has its surface area increased through various processeswhich are more thoroughly described in connection with FIGS. 67 and 68.FIG. 67 is an enlarged, fragmented sectional view of the surface 161taken from FIG. 66. The energy dissipating surface 161 area has beenroughened to create a high surface area, through, for example, plasmaetching 214, chemical etching, or the like. A high surface area can alsobe accomplished by porous coating deposits utilizing physical vapordeposition, chemical vapor deposition or electron beam depositionprocesses. Such porous coating deposits can include fractal coatings,metal nitrides, titanium nitrides, metal oxides, metal carbides, orvirtually anything that would provide a high surface or poroussubstrate. In addition, electrochemical deposition of porous coating,such as iridium-oxide, can also be utilized, as well as nucleate highsurface area morphologically structured coatings, such as whiskers,sub-micron filaments, tubes, nanotubes, or other morphologicalstructures such as columnar, titanium-nitride or iridium-oxide. Any ofthese types of surface conditionings can greatly increase the energydissipating surface area 161. FIG. 68, which is similar to FIG. 67,illustrates the use of carbon nanotubes or fractal coatings 216 toincrease the surface area and therefore the energy dissipation.

FIG. 69 shows a steerable catheter 218, which is typically used for avariety of applications including RF or cryo-ablation, cardiac mappingand many other purposes. Examples of RF ablation include treatment fornephrotic conditions, liver, brain, cancers and the like. For example,this would enable stereotactic ablation of certain lesions within thelung. An emerging field is the entire field of using ablation to treatvarious ventricular arrhythmias, including ventricular tachycardia. Theillustrated catheter 218 in FIG. 69 is meant to be representative of alltypes of catheters or probes which can be inserted into the venoussystem or other areas of the human body. The catheter 218 has a tip 220and an adjacent electrode surface 222, and a main catheter body 224,which can be steered around torturous paths. The steerable catheter 218has a handle 226 which can have various shapes, sizes and configurationsin the prior art. By twisting the illustrated cap 228 of the handle 226,one is able to steer the catheter 218 causing its tip 220 or othersegments to bend as one guides it.

FIG. 70 is an enlarged section taken along line 70-70 in FIG. 69. FIG.70 illustrates that the handle 226 includes an optional but preferredouter insulation sheath 230 which would typically be of plastic orsimilar material that would preferably not be highly thermallyconductive. Inside of the handle (or even the catheter body itself—notshown) 226 are shown in cross-section leadwires 104 and 106. Theillustration of two leadwires is not meant to be limiting since anynumber of wires could be inside the handle 226 and/or catheter 218 tosense electrical activity or deliver ablation energy in accordance withthe present invention, there are frequency selective diverter impedanceelements 112 shown between the leadwires 104, 106 and an energydissipating surface 161, such as a metallic sheath 232. The energydissipating surface 161 does not necessarily have to be metallic, but ithas to be capable of collecting RF energy and conducting thermal energy.This heat energy is therefore dissipated over the large surface area andthermal mass of the catheter body or the handle 226 itself. This resultsin very little temperature rise, but at the same time, accomplishes thegoal of the present invention in redirecting RF energy out of theleadwires 104 and 106 that may be picked up by MRI RF pulsed fields anddirecting said energy into the relatively large surface area 232 insidethe handle 226. Of course, one could eliminate the outer insulationsheath 230. However, in a preferred embodiment, the insulation sheath230 would be relatively poor in thermal conductivity so that one didreally not feel any temperature increase in his or her hand. Referringonce again to FIG. 70, the diverter elements 112 can, of course, becombined with any of the previously mentioned impeder elements such asinductors or bandstop filters.

FIG. 71 is very similar to FIG. 63 except that a number of individual RFenergy or heat dissipating segments 161 ₁, 161 ₂ and 161 _(n) are shown.These are shown spaced apart by separation gaps d₁ and d_(n), which inreality can be quite small. The reason that these energy dissipatingsurfaces are segmented is so that they do not become physically andelectrically long enough to become a significant fraction or multiple ofa wavelength of the MRI pulsed frequency. Such short conductive sectionsdo not pick up significant energy from MRI whereas elongated leadwiresor conductors can, for example, resonate and pick up very significantamounts of MRI RF energy. It would be highly undesirable if the energydissipating surfaces, as illustrated in FIG. 73, were formed to becontinuous along the entire length of the catheter 102 as previouslydescribed in connection with FIG. 71. In this case, the energydissipating surface would actually become an energy collecting surfacebecause it would become a very effective antenna for the MRI pulsed RFsignals. Accordingly, breaking this up into discrete segments preventsthe surface 161 from actually becoming a receiver or antenna for the MRIinduced energy.

FIG. 72 illustrates a paddle electrode 234 which could be used, forexample, in spinal cord simulator applications. It has eight electrodes236 housed in a biocompatible insulative and flexible body 240. Eightleadwires 242 (there can be any number) are connected respectively toeach of the eight electrodes 236. As previously discussed, the elongatedleadwires 242 can pick up significant amounts of RF energy during MRIscanning. It is very important that the electrodes 236 do not overheatsince they are in direct contact with the body, for example, with thespinal cord.

FIG. 73 illustrates the reverse side of the paddle electrode 234, wherean energy dissipating surface 161 is located. As shown in FIG. 74, onecan see that the electrodes 236 are conductive pads that contact thespinal nerve route or at least are closely associated with it. Theleadwires 242 are each electrically connected to respective electrodes236. There is a frequency variable impedance for diverter) element 112in accordance with the present invention shown between each electrode236 and the energy dissipating surface 161. These can be individualdiscrete capacitors or individual discrete L-C traps as shown in FIGS. 5and 6. These can also be one continuous parasitic capacitance elementthat formed between the overlap of each of the electrodes and the areaof the 161 surface itself. In this case, the insulative dielectricmaterial 244 shown in FIG. 74 would be of relatively high dielectricconstant. A high dielectric constant material is desirable so that theamount of parasitic capacitance would be relatively large. By usingparasitic capacitance and appropriate dielectric materials, oneeliminates the need to use individually installed passive circuitelements. Referring to FIGS. 72-74, one can see that the undesirable RFenergy is dissipated on the opposite face of the paddle electrode 234relative to the electrodes that are in contact with the spinal nerveroute. In other words, the RF or thermal energy is dissipated over arelatively large surface area and is directed away from the sensitivejuncture between the electrode body tissue contact area. This isimportant for two reasons, if the RF energy was allowed to concentrateon any one of the electrodes due to resonance phenomenon, then a veryhigh temperature rise could occur which could cause thermal injury tothe spinal nerve itself. By redirecting the energy in the oppositedirection towards the muscle tissue and over a much larger surface area,much less temperature rise occurs, and even if it does, it is directedinto less sensitive tissue.

FIG. 75 illustrates a different type of paddle lead structure 246showing a total of fifteen electrodes 236. In this case there are twoenergy dissipating surfaces 161 and 161. For maximum surface area, theenergy dissipating surfaces could be on the top surface of the paddlelead structure 246, as well as on the backside or back surface (notshown). In accordance with the present invention, FIG. 76 illustrates afrequency selective variable impedance element 112 which is used todivert RF energy from the electrodes 236 to the 161 surfaces.

FIG. 77 is very similar to FIGS. 31, 32 and 33 in that it shows asection of human head with a deep brain stimulator disposed therein.There are a plurality of leadwires 104 and 106 which are connected to anAIMD or pulse generator (not shown). The pulse generator would typicallybe placed in the pectoral region and leadwires 104 and 106 would berouted up along the patient's neck to the deep brain electrodes 108 and110. Referring to FIGS. 77-79, one can see that there is a novel tether248 or wire arrangement where the leadwires 104, 106 are not onlyconnected to the distal electrodes 108, 110, but they are also connectedto a pair of energy dissipating surfaces 161 and 161. In FIG. 78, onecan see the tether area 248 wherein the leadwires 104, 106 connectindividually to the electrodes. As shown in FIG. 79, the leadwires 104,106 have a connection inside the tether area 248 such that the wires arerouted both to the distal electrodes 108 and 110 and also throughrespective junctions 250 a and 250 b to two individual energydissipating surfaces (161 and 161). The leadwire 104 has a directelectrical connection at junction 250 a to distal electrode 110. Inturn, leadwire 106 has a direct connection at junction 250 b to distalelectrode 108. However, at the junctions 250 a and 250 b, also connectedare frequency selective elements 112 which in turn are connected torespective energy dissipating pad or surfaces 161 and 161′. Of coursethe separate energy dissipating pads could be one large energydissipating pad. However, in order to maximize surface area andfacilitate surgical implantation, two pads are shown. These areoriginally implanted by the physician underneath a skin flap which isthen sewn back down in place. In this way, any heat that is generatedduring MRI procedures is generated on the top side of the skull wellaway from any brain matter.

It will be obvious to those skilled in the art that the presentinvention can be extended to a number of other types of implantablemedical devices, including deep brain stimulators, spinal cordstimulators, urinary incontinence stimulators and many other types ofdevices.

FIG. 80 is an overall outline drawing showing a cardiac pacemaker 102with endocardial leads LW_(1,2) and LW_(3,4) implanted into a humanheart 136 as shown. Each lead is bipolar meaning that it contains twoleadwires 104 ₁, 104 ₂ and 106 ₃, 106 ₄, One can see that lead 104 ₁,104 ₂ is routed into the right atrium and that leadwire 106 ₃, 106 ₄ isrouted into the right ventricular apex (RV). The distal electrodes forthe atrial lead are shown at tip 144 and ring electrode 108. In theright ventricle, the distal electrode tip 138 is shown in closeproximity to distal ring electrode 110. As previously mentioned,bandstop filters in accordance with U.S. Pat. No. 7,363,090 could heplaced at or near the distal electrodes 138, 110, 144, 108 as needed.Referring to the AIMD housing, one can see that there are variableimpedance elements 112 and 118 associated with each one of the leadwires104 ₁, 104 ₂ and 106 ₃, 106 ₄.

FIG. 81 is an outline drawing of an AIMD such as a cardiac pacemaker.Shown is a metallic, typically titanium, housing 124. Its housing 124hermetically sealed with a laser weld 176 as shown. It has a hermeticseal 128, which is also laser welded into the titanium housing 124. Thehermetic seal has an insulator, which is well known in the prior art,through which leadwires 104 ₁, 104 ₂ and 106 ₃, 106 ₄ pass through innon-conductive relationship with conductive housing 124. A typicalpacemaker connector block 134 is shown. This can be in accordance withvarious international standards organization (ISO) such as IS-1, DF-1,IS-4 and the like. Connector ports 130 allow for convenient connectionof a lead, which can be routed to the appropriate body tissue to besensed or stimulated. Referring once again to FIG. 83, one can see thatthe leadwires 104 ₁ through 106 ₄ are generally routed to circuit boards(f), integrated circuits or substrates within the active implantablemedical device housing 124. These can include cardiac sense circuits,pace circuits and the like. Referring once again to FIG. 81, one can seethat there are variable frequency impedance elements 112 and 118 asillustrated on leadwire 106 ₄. It should be noted that these variablefrequency impedance circuit elements would appear on all or some of theleadwires 104 ₁ through 106 ₄. They are only shown on 106 ₄ to simplifythe drawing. In this example, the metallic housing (titanium) 124 of theAIMD as an energy dissipating surface 161. Typically the AIMD isinstalled in a pectoral pocket, an abdominal pocket or in some otherlocation that is not in intimate contact with a body organ. Accordingly,if the housing 124 were to overheat, it would be surrounded by fat andmuscular tissue which is not nearly as sensitive to thermal damage as,for example, cardiac tissue or brain tissue. Also referring back to FIG.81, one can see that for AIMDs, the relative surface area of the housing124 is quite large in comparison to the electrode at or near the end ofan implanted lead. In other words, it embodies a great deal of surfacearea over which to dissipate the MRI RF energy. Accordingly, the thermalrise will be very low (just a few degrees) as opposed to if the energywere concentrated over a small area in electrode tip where the thermalrise can exceed 30 or even degrees centigrade. Accordingly, it is aprimary feature of the present invention that the housing of the AIMD beused as an energy dissipating surface 161 working in combination with orwithout bandstop filters installed at or near the distal electrode totissue interface. In FIG. 81, this energy dissipation is represented bythe arrow marked 161. In fact, the energy is being dissipated at allpoints all around the metallic housing 124 to the surrounding tissues.

Referring once again to FIG. 81, the diverter element 112 can be any ofthe diverter elements 112 described in FIGS. 4-11 herein. Impededelement 118 can be any of the impeded elements 118 that are alsoillustrated in FIGS. 4-11 herein. The impeder and diverter elements 112and 118 can also be placed within a molded header block 134 of a cardiacpacemaker or the like. These are shown as diverter element 112A andimpeder element 118A. The impeder elements 112 and 118, as illustratedin FIG. 81, are desirably but not necessarily used in combination. Inother words, only the diverter element 112A could be used. An advantageto locating the diverter element 112A and/or impeder element 118A in theheader block 134 (outside of the AIMD housing 124) is that the componentvalues of these impeder and diverter elements could be optimized tomatch a particular lead. Typically, cardiac pacemakers are manufacturedin high volume (there are over 600,000 pacemakers manufactured everyyear). It is really not practical to build a custom pacemaker for everytype of lead. Leads vary in length from about 20 centimeters to over 60centimeters, depending on whether its a pediatric or a large adultapplication. Accordingly, the characteristic impedance of the lead willvary widely with its length and also its implanted lead trajectory inhuman tissues. By having the diverter and/or impeder elements 112A and118A located in the header block, then they can be easily customized tomatch a particular lead. In a particularly preferred embodiment, thediverter element 112 and/or the impeder element could also be placed inthe proximal end of the lead or even in the male proximal leadconnector. In this way, each lead could be optimally tuned with adiverter such that maximal energy transfer would occur between theimplanted lead and the medical device housing 124. Referring once againto FIG. 81, diverter element 112B and impeder element 118B are showndisposed within a proximal lead connector. In accordance with thepresent invention, in order to maximize energy transfer, the diverterelement 112B would be tuned and opposite to the characteristic impedanceof the lead. These principles are more fully described in FIGS. 90-94.

FIG. 82 illustrates that variable frequency impedance diverter element112 can be any type of capacitor element, including MLCC chip capacitorsand the like. FIG. 84 illustrates that the variable frequency impedanceelement 112 can also be a feedthrough capacitor as has been noted is inthe prior art. Referring once again to FIG. 83, one can see there is aresistor 260 in series with the capacitor element 114. This resistor cansimply be the equivalent series resistance (ESR) of the capacitor 114itself, or it can be a separate discrete resistor mounted as a separatecomponent. Ideally, the value of the resistance 260 would be equal tothe characteristic impedance of the implanted lead at the same time thecapacitor's, 114 capacitive reactance would tend to cancel out theinductive reactance of the implanted lead. This would facilitatetransfer of a substantial amount of RF energy out of the leads to thesurface 161. It will be obvious to those skilled in the art that adiscrete resistor 260 could be added to any of the diverting elements ofthe present invention including diverting elements such as illustratedin FIG. 83. Resistors can also be associated with any of the low passfilters as previously described in FIG. 43. For a feedthrough capacitoras illustrated in FIG. 84, it is difficult to control the amount ofseries resistance (ESR) to achieve optimal energy transfer to the 161surface. One is referred to U.S. Pat. No. 7,623,336, the contents ofwhich are incorporated herein. FIGS. 3 through 6 illustrate novel waysto deliberately increase the feedthrough capacitors equivalent seriesresistance (ESR). These are methods to increase the resistivity of theelectrode plates at high frequency. The capacitor ESR can also bedeliberately increased by using fewer electrodes (this necessitatesthinner dielectric) and also deposition of thinner electrode plates.There is a tradeoff here between the power handling ability of thefeedthrough capacitor. Accordingly, the preferred embodiment that adiscrete chip resistor and capacitor would be used as illustrated inFIG. 83.

FIG. 84 is a close-up view of the variable impedance elements 112 and118 from FIG. 81 located within the housing of an AIMD can 124. Aspreviously mentioned, the variable impedance elements 112 and 118 wouldbe installed on all of the leads that ingress and egress the AIMD. Theground symbol 256 is shown to indicate that variable impedance element112 is connected to the housing 124 of the AIMD. The leadwire lengthsshould be very short to minimize inductance. These sections of leadwire258 ₁ and 258 ₂ are kept very short so that high frequency energy fromMRI will, not be reradiated to sensitive AIMD circuits. Ideally, circuitelement 112 would be a chip which would be bonded right at the point ofleadwire ingress and egress.

FIG. 85 illustrates that the trap filter of FIG. 45 can be used incombination with a second diverter such as a capacitor 114 as previouslyillustrated in FIG. 83 or a feedthrough capacitor as illustrated in FIG.84. For a pacemaker or an 100, this would be the most common embodiment.Typical capacitance value for the series resonant trap would be 270nanohenries of inductance and 22 picofarads of capacitance. This wouldmake the series trap filter series resonant at 64 MHz. Its alsoimportant that the designer realize that at a certain frequency, thecombination of the trap filter 206 and the EMI filter C_(X) will at somepoint become a parallel resonant bandstop filter. This happens atfrequencies at which the trap filter becomes inductive. In other words,at resonance, the inductive reactance cancels out the capacitivereactance and the impedance of the series trap is essentially zeroexcept for its real or resistive losses. As previously mentioned,ideally the value of resistor 260 is selected such as to be equal to theequivalent or characteristic series resistance of the implanted leadsystem for maximal energy transfer. However, at frequencies aboveresonance, the inductive reactance term tends to increase and dominatethe capacitive reactance term. In other words, at frequencies aboveresonance the series L-C trap will, tend to look like an inductor whichcould then cause a secondary resonance in parallel with the capacitor114 _(x). This means that there would be a minor degradation in theoverall attenuation to electromagnetic interference. This resonant pointshould not appear at the frequency of a new and powerful emitter.Resonance at these emitter frequencies therefore should be avoided.Bandstop filter BSF′ is optional, but does separate the divertercapacitor CX from the L-C trap filter.

FIG. 86 is essentially the same as FIG. 82 except the focus is on theseries variable impedance impeder element 118. The use of a seriesimpedance element 118 is optional, but highly desirable for AIMDs thathave sense circuits.

FIG. 87 indicates that the variable impedance impeder element 118 can bean inductor 116 as shown. This forms what is known in the art as asingle element low pass filter. The inductor element 116 would freelypass low frequencies such as biologic frequencies that would offer ahigher impedance at high frequencies such as those of MRI pulsefrequencies, cellular telephones and the like.

FIG. 88 illustrates that the variable impedance element 118 can be aparallel resonant L-C bandstop filter BSF as shown. The operation of thebandstop filter has been clearly described in U.S. Pat. No. 7,363,090and US 2007/0112398 A1.

FIG. 89 shows a unipolar lead system 104.sub.1 for an active implantablemedical device. A unipolar lead system is shown for simplicity. It willbe obvious to those skilled in the art that any number of lead wires 104₁ could be used. In FIG. 90, one will see that this system involves anAIND and its associated housing 124 attached to unipolar lead wire 104 ₁to a human heart 136. At the distal tip or distal end of lead wire104.sub.1 is an optional bandstop filter 117. The optional bandstopfilter 117, which is located at or near the distal electrode 138, ismore thoroughly described in. U.S. Pat. No. 7,363,090 the contents ofwhich are incorporated herein. As shown, the implanted lead hasinductive 116 and resistive 260 properties along its length (it may alsohave capacitive properties as well). The total or equivalent, inductivereactance of the lead in ohms is given by the formula as shown in FIG.89. As mentioned, the distal electrode handstop filter 117 may or maynot be present. The equivalent inductance 116 and resistance 260 of thelead system also includes the impedance of any tissue return path. Itshould be noted that the present invention applies to any type of AIMDincluding those AIMDs whose housing may actually be an active electrodeor part of an electrode return path. For example, there are certainneurostimulator applications involving a number of distal electrodesthat all have return paths through body tissue from a digital electrode138 all the way to a common electrode which is also the device housing.One of the best ways to actually determine the characteristic leadimpedance, including its inductive, capacitive, and resistiveproperties, is through human body modeling using software such asSAMCAD, Using SAMCAD, one can calculate the electric field vectors allalong the lead trajectory. One can then calculate the induced energyinto the implanted leads and their characteristic impedances. Referringonce again to FIG. 89 one can see that on the interior of the generallymetallic housing of the AIMD 124 there are frequency selectivecomponents 112 and 118. These frequency selective elements can consistof various arrangements of capacitors, inductors and resistors or evenshort circuits as will be more fully described in FIGS. 90 though 93.

FIG. 90 illustrates the lead system of FIG. 89 wherein an L-C trapfilter 206 has been placed at the point of leadwire ingress into thehousing 124 of the AIMD. In this case, L_(s) and C_(s) have beendesigned to be resonant at the pulsed RF frequency of the MRI equipment.Therefore, this forms an RF short to the AIMD housing 124 which becomesan energy dissipating surface 161 of the present invention. Aspreviously described, a series resistance could be added in series with116, and 114, in order to further optimize energy transfer from theleadwire 104. It is desirable that this surface area be relatively highso that very little temperature rise occurs on surface 124 as the MRI RFenergy is being dissipated.

FIG. 91 is another illustration of the unipolar lead system of FIG. 89.In this case, diverter element 112 features a capacitive element 114whose capacitive reactance is given by the equation −j/ωC. In apreferred embodiment, the inductance of the implanted lead would firstbe modeled, calculated or measured. Therefore, the value of capacitancecould be tuned or selected such that −j/ωC is equal and opposite to+jωL. In this case, the reactances cancel each other so that one getsmaximal energy transfer to the energy dissipating surface 124, 161. Aspreviously described, the capacitor's equivalent series resistance (ESR)could be controlled or a discrete resistance approximately equal to thecharacteristic resistance of the implanted lead could be added in seriesin order to further maximize energy transfer from the implanted leadsystem 104 ₁ to the 161 surface 124.

FIG. 92 embodies the simplest arrangement wherein lead wire 104.sub.1 issimply shorted to the housing 124 of the AIMD which makes said housingan efficient energy dissipating surface 161. Of course, creating a shortto housing as illustrated in FIG. 92 would also short out the properoperation of the AIMD. This generally would not be acceptable for alifesaving device such as cardiac pacemaker. However, for aneurostimulator such as a spinal cord pain control stimulator, thiscould be a programmable function wherein the AIMD leads were shorted outonly for the MRI scan.

FIG. 93 is similar to the unipolar lead system previously described inFIGS. 89 and 91. In this case, as for FIG. 91, the capacitance value 114has been selected such that the capacitive reactance will be ideallyequal and opposite to the inductive reactance of the implanted lead.However, in this case, the resistances are also balanced. In otherwords, the resistance of the implanted lead 260 is equal in value to adiscrete resistor 260 placed inside or outside of the housing 124 of theAIMD. In this case, maximum power transfer or energy will be dissipatedby this internal resistance 260 _(X) as heat. In a preferred embodiment,a thermally conductive but electrically insulative filler material (notshown) will be placed between the resistor 260.sub.X and the AIMDhousing 124 such that maximum energy transfer from resistor 260 _(X)will occur. In fact, in a preferred embodiment, resistor 260 _(X) shallhave a finned high surface area housing for maximal energy transfer tothe surrounding encapsulant. Referring once again to FIG. 93, one cansee that energy is radiated and conducted from a discrete resistanceelement 260 _(X) shown as 161. This energy being dissipated turns tothermal (heat) energy. It is desirable to have a relatively largethermal mass located within housing 124. The AIMD housing 124 thenbecomes a heat dissipating surface 262. This thermal energy will bedissipated over the relatively large surface area 124 into body fluidsand tissues that surround the AIMD. For example, in a cardiac pacemakerapplication, housing 124 would be in a pectoral muscle pocket.

Referring back to FIGS. 91 and 93, it is not necessary that thereactances completely cancel, or in the case of FIG. 93, it's notparticularly important that the resistances are exactly equal. In fact,there is a tradeoff between EMI filtering of the input capacitance andexact cancellation of the +jωL component lead system. As it turns out,through actual testing, it is really only important that the impedancegenerally be cancelled in the lead system so that at least the bulk ofthe excess energy from the MRI RF pulse field will be dissipated to thehousing of the AIMD 124. For example, if one calculates that a 75picofarad capacitor would exactly cancel the inductive reactance of thelead system, one may instead choose to use a 1000 picofarad capacitor.The 1000 picofarad capacitor would still draw a large amount of energyfrom the lead system to the housing 124. The reason one would do this,is that a 1000 picofarad capacitor would offer much more effective EMIfiltering to not only the RF pulse frequency (64 MHz or 1.4 Telsa MRsystem), but also for cell phones and other emitters commonly found inthe pace environment. The energy balance systems and circuits of thepresent invention can also be combined with knowledge of the implantedlead design. By varying the coil pitch on the outer coil to createenough impedance so that a pacemaker ring electrode does not heat up ispossible when combined with a bandstop filter and the distal tipelectrode circuit. This becomes a balancing act so that one can be surethat not too much energy is transferred from the ring electrode to theinner electrode lead coils.

FIG. 94 is a schematic diagram of an implanted lead system very similarto that which was previously described in FIG. 89. Shown is an implantedlead 104.sub.1 which is directed into body fluids and to at least onedistal electrode. Shown is a diverter element 112 in accordance with thepresent invention and an optional impeder element 118 also in accordancewith the present invention. The novel switch 252 shown in the openposition. The switch 252 is inclusive of all types of mechanical orelectronic or microelectronic switches. This includes mechanicalswitches, DIP switches, MEMS switches, microelectronic switches,microelectronic switch arrays, any type of electronic switches includingfield effect transistor (FET) switches, varistor type switches and thelike. In a preferred embodiment, switch 252 is programmable through AIMDtelemetry. In a probe or catheter application, a signal can be sent intothe probe or catheter to cause the switch 252 to switch positions. Inanother preferred embodiment, switch 252 could be automaticallyactivated by a static field sensor. As previously described, there arethree main fields associated with magnetic resonance imaging. This isthe main static B₀ field, which for most modern scanners is either 1.5or 3.0 Tesla, There is also an RF field and a gradient field. The switch252 as illustrated in FIG. 94 can be associated with a static field B₀sensor. As shown in FIG. 98, this could be a Hall effect device, a reedswitch, a ferrite chip or any other type of magnetic field sensor. Inthis way, no device reprogramming would be necessary. In other words,when the patient is introduced to the MRI bore, the B₀ field sensor SFSlocated in the AIMD and/or the probe, catheter or the like inassociation with switch 252 would automatically sense the presence ofthe MRI main static field, thereby switching the switch 252 into theclosed position (the switch shown in FIG. 94 is shown in the openposition). The reason for the configuration as illustrated in FIG. 94 isthat in order to provide for optimal energy transfer, the frequencydiverting element 112 may place a burden (electrical load) on the AIMDduring normal operations. In other words, it may degrade pacing pulsesor rob energy during normal operation. Using the novel circuit asillustrated in FIG. 94, the frequency diverting element 112 could beconductively coupled to the energy dissipating surface only when neededduring MRI scans. The switch diverter of FIG. 94 is for minimizingheating of an implanted lead and/or its distal electrodes when in thepresence of a high power electromagnetic field environment such as thatcreated by the RF pulse field of an MRI scanner. The diversion circuit,as illustrated in FIG. 94, can even embody a short circuit as will befurther described. The frequency diverting element 112 can also be abandstop filter as previously described as element 117 in FIG. 11. Sincethe novel switch 252 is only closed during exposure to highelectromagnetic field environments, this enables a wider variety ofimpeder elements 112. It is also important to note that this also opensup a wider variety of passive component values for element 112. Forexample, when diverter element 112 is a capacitor as described in FIG.5, it can now be of a very high capacitance value. This may rob someenergy from the AIMD and even cause some miniscule battery depletion;however, this is unimportant during the relatively short time periodrequired to complete an MRI scan compared to the overall lifetime of theAIMD.

FIG. 95 is very similar to FIG. 94 except that the novel switch 252 isshown disposed between the implanted lead or leadwires or wiring insideof the AIMD in diverter circuit 112.

Since the diverter is in series, it really does not matter where theswitch 252 appears. Referring once again to FIG. 95, one can see thatthe circuits described can also be associated with an optional impederelement 118 that has been previously described in FIGS. 7, 10 and 11.

FIG. 96 illustrates that two switches 252 and 252′ may be employed.Switch 252 is shown in the open position and works in a very similarfashion to switch 252 that was previously described in FIGS. 94 and 95.In this case, the diverter element 112 is shown as a short circuit tothe housing. This would not be desirable for a cardiac pacemaker patientwho was pacemaker dependent. In this case, the patient depends on thebeat pulse or the pacemaker to provide each and every heartbeat. Itwould be disastrous for each and every patient to short out their deviceto the energy dissipating surface 161. Referring to FIG. 96, the switch252 is shown open which would be the normal operating mode. Inpreparation for and during an MRI scan, switch 252 would be switched tothe closed position which would short an associated leadwire 104.sub.1or implanted lead to the energy dissipating surface 161. This would workfor many types of AIMDs including neurostimulators and the like, wherethe therapy they provide is convenient, but not life-sustaining.Accordingly, for the MRI compatible mode as shown in FIG. 97, switch 252would be closed, thereby shorting or diverting energy from leadwire104.sub.1 directly to the energy dissipation surface 161. As previouslydescribed, the short could be replaced by any of the diverter circuits112 as previously described. Referring once again to FIG. 96, one cansee that there is a second switch 252′. This switch would normally beclosed, as shown, during normal AIMD, probe or catheter operations.However, this switch 252 would be opened as shown in FIG. 97 during MRIscanning. This has a desired effect of disconnecting the AIMDselectronic circuits from the implanted lead system during MR scans. Thisis very important to provide a very high level of EMI protection to AIMDcircuits and also to prevent an effect known as RF rectification. MRI RFpulses consist of high frequencies that are determined by the Lamourfrequency. For a 1.5 Tesla scanner, this means that there would be 64MHz bursts of RF energy. These packets of energy, if detected by anon-linear element (like a diode) in the AIMD, could turn into theequivalent of digital pulse trains. This has the potential to directlycapture the heart, for example, at 200 beats per minute or higher,thereby introducing a very dangerous tachyarrhythmia. Ventriculardefibrillation, for example, can be immediately life-threatening. Bydisconnecting AIMD electronics during MRI scans, one eliminates anypossibility of RF rectification. A discussion of where non-linearelements come from in an AIMD is important. Almost all modern AIMDs haveover-voltage circuit protection. There typically are in the form ofdiodes, Transorbs, or the like. These diodes act as detection circuitsand can strip off the pulse modulation from MRI RF frequencies.Therefore, its important that switch 252′ be opened up prior toencountering any protection diodes or non-linear elements inside of theAIMD. Pacemaker or neurologerlator sense circuit can also become verynon-linear in the presence of high amplitude signals. This has to dowith limitations of the dynamic range of their electronic (active)bandpass of other filters or amplifiers. There is also another problemthat can be associated with MRI scans and that is induced currents onthe implanted leads to the MRI gradient fields. In general, MRI gradientfields are low frequency and are generally in the 1-2 kHz (maximum 10kHz) frequency range. At low frequencies, such as MRI gradientfrequencies, currents are induced in implanted leads primarily byFaraday's Law of Induction. This is further explained herein in FIG.101. Since there is a loop current involved that flows from the leadsthrough the AIMD and then from the AIMD case back through body tissues,it is very useful if the AIMD has a very high impedance. The highimpedance at the AIMD limits the amount of current flowing in this loop.There is also a problem associated with what is known in the industry asgradient induced rectification. This is very similar to RF rectificationin that low frequency gradient currents consist of pulses. If thesepulses encounter an AIMD or catheter non-linear circuit element, such asa diode, they can become demodulated, wherein they end up showing up onthe leads or in the implanted medical device as a series of lowfrequency pulses. This can have an EMI effect, for example, inhibit apulse generator which might falsely interpret these gradient pulses as anormal heart beat. Worse yet, these gradient pulses can appear on theimplanted lead as a dangerously high biologic frequency (for example, 50Hz) where the pulses would directly capture the heart and inducefibrillation. For a neurostimulator, they can also directly induce pain,for example, in the spinal cord. By having switch 252 open, asillustrated in FIG. 97, one disconnects AIMD internal electroniccircuitry. In a preferred embodiment, the switch 252′ would disconnectall non-linear circuit elements including AIMD high voltage protectiondiodes, bandstop filters, non-linear active filters and the like. It isalso very important that the switch 252 itself not become non-linear.This would definitely be a possibility for certain types of electronicswitches that were not robust enough. Accordingly, it is a principle ofthe present invention that switch 252′ be of a type that will not becomenon-linear in the presence of MRI gradient or RF current/voltages.Referring once again to FIG. 94, a short circuit to the energydissipating surface housing would not be the preferred embodiment forelongated lead exposure to gradient fields. A preferred embodiment woulduse a frequency selective diverter 112 as is shown in FIG. 95 and asdescribed in FIGS. 2-11. The frequency selective diverter 112 acts as ahigh pass filter, meaning that it will allow 64 MHz or other MRI RFfrequencies to flow freely to the EDS housing. However, the diverterelement 112 will look like a very high impedance or open circuit atextremely low frequencies, such as MRI gradient frequencies. This tendsto open the loop thereby preventing excessive currents from flowing inthe implanted lead and through body tissues, such as cardiac tissue. Ina particularly preferred embodiment, the diverter 112 would be acapacitor as shown in FIG. 91 or an L-C trap filter as shown in FIG. 90or a resistor in series with a capacitor as illustrated in FIG. 93.

Referring once again to FIGS. 94, 95, 96, and 97, it is not necessary oreven required that the switches 252 or 252′ be installed in each andevery AIMD implanted lead circuit. For example, for an implantablecardioverter defibrillator (ICD), there are typically sense circuits,high voltage shock circuits and pacing circuits. One can install thenovel switches 252 and 252′ only as required to achieve optimalperformance. For example, referring once again to the ICD, one maychoose to leave a pacing circuit in its normal operating mode whileswitching all other AIMD implanted leads to an MRI compatible mode.Referring once again to FIG. 96, switch 252 can appear on the right handside of Node 264 as shown or it could be placed on the left hand side ofNode 264 (not shown).

FIG. 98 is very similar to FIGS. 96 and 97 except that switches 252 and252′ have been replaced by a single switch 266 ₁ which can be either asingle or multipole double throw switch as illustrated. In other words,the switches as previously illustrated in FIGS. 96 and 97 have beencombined into a single switch. FIG. 99 is shown for lust one circuit ofan AIMD or a probe or a catheter. It will be obvious to those skilled inthe art that any number of switches 266 ₁ can be installed in any or allof the implanted leads as required. The switch 266 ₁ as shown in FIG. 98can be disposed anywhere along the implanted lead or inside of an energydissipating surface 161 or inside of an AIMD housing 124 as shown, oreven anywhere along the handle or wiring or body of a probe or catheter.Switch 266 ₁ could also be installed in an external electronics box, forexample, that connected to the handle of a probe or catheter, a looprecorder or the like.

Referring once again to FIG. 98, one can see that the switch is shown inposition X which connects the leadwire to the diverter circuit 112. Aspreviously mentioned, diverter circuit 112 can be any of the frequencyselective diverter circuits, a short circuit or even a bandstop filter.When switch 266.sub.1 is thrown to position X as shown, it is in the MRImode. During normal operations, it would be switched to position Y sothat the AIMD, probe or catheter could normally operate. Any of theswitch technologies previously described in FIG. 94 apply to any and allof the switches described herein.

FIG. 99 illustrates that the switch 266 ₁, as previously described inFIG. 98, can also be disposed within an energy dissipating surface 161located anywhere along an implanted lead. The use of switches, asillustrated in FIGS. 94 through 99, are applicable to any of the designsillustrated herein.

FIG. 100 shows a preferred embodiment of the switch 266 ₁ and diverter112 that was previously described in FIG. 98. Shown is a capacitor 114that is connected between switch point X and the energy dissipatingsurface 161, such as the housing of an AIND, or body of a probe orcatheter. Capacitor 114 as shown or a resistor in series with acapacitor or an L-C trap filter (not shown) are preferred because theytend to look like high pass filters in an MRI environment. MRI equipmentproduces three main fields consisting of the main static field B₀, theRF pulsed field, and the gradient field. We can ignore the static fieldfor these purposes since it does not induce currents on implantedleadwires. It is a basic principle of physics that either the field hasto be moving or a conductor has to be moving in relation to the magneticfield for there to be an induced electromotive force. Both the RF pulsedfield and the gradient field are time varying and can induce undesirablecurrents on implanted leads. Capacitor 114, as shown in FIG. 100, is afrequency variable impedance diverter element in accordance with thepresent invention. Accordingly, it has a very high impedance at lowfrequency and a very low impedance at high frequency (hence, when wiredin a circuit like this, it is known as a high pass filter). It,therefore, shunts or diverts undesirable high frequency RF energy to theenergy dissipating surface housing while at the same time blocking theflow of low frequency energy such as that induced by MRI gradientfields. Preventing the flow of gradient currents into AIMD circuitry isvery important so that gradient rectification does not occur. Inaddition, opening up the loop at gradient frequencies, as shown in FIG.100 (effectively open by the high impedance of the capacitor 114), alsoprevents flow of current through distal tissues at theelectrode-to-tissue interface.

FIG. 101 is a patient front view representative of an X-ray tracing ofan implanted cardiac pacemaker. The pacemaker is shown installed in apectoral pocket, which can either be left pectoral (as shown) or rightpectoral (not shown). There can be one or more turns of excess leadwirethat's coiled up in the pectoral pocket and then the lead is routedendocardially down through the superior vena cava into cardiac chambersas shown. A loop area shown by the checker pattern is formed between thedistal tip electrode all along the lead to the AIMD housing and thenthrough a multi-path tissue return path shown as a dashed line from theAIMD housing to the distal tip electrode in the heart. MRI low frequencygradient fields couple into this loop by Faraday's Law of Induction. Ingeneral, Faraday's Law states that a voltage induced in this loop isdirectly proportionate to the area of the loop times the rate of changeof the magnetic field in Teslas per second. A worse case couplingsituation occurs when the field is orthogonal to the loop area. Currentwill flow in the lead unless the lead is opened up (switched open). Itis highly undesirable that this low frequency MRI gradient inducedcurrent flow into cardiac tissues as this could directly induce cardiacarrhythmias. It is also undesirable if this current should flow into theAIMD electronics as it could either interfere with AIMD electronics(EMI) or it could lead to gradient rectification. In the art, directcardiac or tissue stimulation is known as Gradient STIM.

FIG. 102 is the high frequency model of the circuit illustrated in FIG.100. In this case, at high frequencies, such as MRI RF pulsedfrequencies, the capacitor 114 is a very low impedance which effectivelyappears as a short circuit. This has the desirable effect of pulling ordiverting high frequency energy on the lead 104.sub.1 through the lowimpedance of the capacitor 114 to the energy dissipating surface 161,which in this case, is the AIMD housing. As previously stated, whenwired between point X and 161, the capacitor 114 acts as a high passfilter. The capacitive reactance of capacitor 114 as previouslydescribed is a −j vector which tends to cancel the +j vector that isassociated with the inductance of an implanted lead. As previouslydescribed, this aids in maximal (tuned) energy transfer to the energydissipating surface 161.

FIG. 103 is the low frequency model of the circuit previouslyillustrated in FIG. 100. In this case, at low frequencies, the capacitor114 appears as a very high impedance which effectively appearselectrically as an open circuit. As previously mentioned, this has thedesirable effect of preventing gradient currents from flowing in theimplanted lead and the associated loop through body tissue and AIMDelectronics.

FIG. 104 shows the application of the switch 266 ₁ of FIG. 98 to acircuit, board 137 or housing located inside of the housing 124 of anAIMD. Shown is a hermetic terminal 128 that was previously described fora cardiac pacemaker in FIG. 12. The overall AIMD housing 124 is shownconnected to the ferrule 129 of the hermetic terminal 128. Leadwires 104_(1,2) through 106 _(3,4) pass through the hermetic terminal innon-conductive relation as previously described in FIG. 12 (only 104 ₁is shown). Internal leadwires are routed from the hermetic terminal 128through circuit terminals 272 to the circuit board or substrate 137.Shown is an optional shielded conduit 268 of the grounded 256 and 256 toa ground trace located within the wiring or flex cable. The switches 266₁ through 266 ₅ could, of course, all be incorporated into a singlemicroelectronic chip. In the art, the switch 266 ₁ would be known as amultipolar double throw switch. Switching the ground circuit 270 would,of course, be optional. As previously described, one could switch any orall of the implanted lead circuits 104 _(1,2) through 106 _(3,4) asshown. In this case, the implanted leads (not shown) and associatedleadwiring 104 _(1,2) through 106 _(3,4) are all shown shorted tocircuit ground 256.

FIG. 105 is very similar to FIG. 104 except that the short has beenreplaced by a frequency variable impedance element 112 in accordancewith the present invention. As previously described in FIG. 2, variablefrequency element 112 could be a capacitor 114, could be an L-C trapfilter consisting of a capacitor 114 in series with an inductor element116, or even a short or a bandstop filter as previously described.

FIG. 106 is very similar to FIGS. 104 and 105 except that the diverterelement 112 is shown as a capacitor 114.

FIG. 107 is very similar to FIG. 106 except that a resistor has beenadded in series with the diverter capacitor 114 as shown. One isreferred to the drawing description for FIG. 93 to see why adding aresistance in series with a capacitor can lead to tuned energy balancein order to draw the maximum amount of induced energy out of the leadsinto the energy dissipating surface 161.

FIG. 108 is very similar to FIGS. 104-107 except that the diverterelement 112 is an R-L-C trap filter. Resistance has been added tocontrol the Q and also tuned for maximum energy transfer in accordancewith FIGS. 45 and 46.

FIG. 109 illustrates one method of connecting a flex cable 274 to leads104 _(1,2) through 106 _(3,4) that pass through the hermetic terminal128 in non-conductive relationship. The conductive ferrule 129 as shown,which would be laser welded into the conductive housing 124 of the AIMD.Leadwires 104 _(1,2) through 106 _(3,4) are routed through the flexcable for convenient attachment to a circuit board or substrate locatedwithin the AIMD.

FIG. 110 is a cross-sectional view taken generally from the attachmentof the flex cable 274 to the hermetic terminal 128 previously describedin FIG. 109. Shown are optional ground shield traces 276 and 276′ whichserve two purposes. The first is to prevent re-radiation of EMI insideof the AIMD housing. The second is to provide a very long impedanceground connection from the AIMD housing to the circuit board orsubstrate where the novel switches of the present invention arepreferably located. Referring once again to FIG. 110, one can seecircuit trace 104 which representative of circuit traces 104 _(1,2)through 106 _(3,4). On the left side, there is a ground connection 256to the ferrule 129 of the hermetic terminal 128. Ground shield traces276 and 276′ sandwich or surround the flex cable or substrate circuittraces 104. Desirably, parasitic capacitances 114 _(P) are formedbetween the circuit traces 104 and the surrounding ground planes 276,276′. The parasitic capacitance 114 _(P) acts as additional low passfiltering.

FIG. 111 is taken generally from section 111-111 from FIG. 110. Thisillustrates the internal circuit traces 104 _(1,2) through 106 _(3,4).Also shown are a number of via holes 278 which are used to stitchtogether and electrically reduce the impedance of the internal groundshield plates 276 and 276′.

FIG. 112 is one of a pair of coaxially surrounding shields disposedabout the circuit traces 104 in non-conductive relation and is takengenerally from section 112-112 from FIG. 110 and shows a preferred formof an upper and lower ground plane 276, 276′. As mentioned, thiseffectively shields circuit traces 104 _(1,2) through 106 _(3,4) andalso at the same time provides a very low impedance RF ground connectionto a switch or switch network located on a remote AIMD circuit board orsubstrate 137. It will be obvious to those skilled in the art that thepair of coaxial shields that are shown in FIG. 112 could also be atubular shield or shielded conduit that extends from the hermeticterminal to the circuit board and provides the same function. This formsa shielded conduit which prevents EMI re-radiation from the leadwiresinside of an AIMD and at the same time, extends a low impedance RFground to the circuit board or substrate. This is useful, not just forthe novel switches and impeder and diverter elements of the presentinvention, but is also very useful to facilitate on board EMI filtering,high voltage suppression arrays, and the like. All of these depend on alow impedance RF ground to the AIMD housing.

FIG. 113 is taken generally from section 113-113 of FIG. 110 and showsan alternative to the layer previously described in FIG. 111 in that thecircuit traces 104 _(1,2) through 104 _(3,4) are shown furthersurrounded by grounds 256 as shown.

FIG. 114 illustrates the flex cable arrangement 274 previously describedin FIGS. 109 through 113 connected to a circuit board or substrate 137which is generally located inside an AIMD housing, inside a probe orcatheter body, or inside of an energy dissipating surface 161. A flexcable 274 is shown connected to circuit traces on the circuit board orsubstrate. Also shown is an electronic switch module 280 in accordancewith the present invention. This electronic switch module could be adiscrete module as shown in FIG. 114, or it could be integrated withinan overall microelectronic chip on the AIMD. Preferred schematicdiagrams for the arrangements illustrated in FIG. 114 have beenpreviously described in FIGS. 104 through 108.

FIG. 115 is very similar to FIG. 81 except that it shows optionallocations for the frequency selective diverter elements 112 and thefrequency selective impeder elements 118 of the present invention. Asshown in FIG. 115, impeder elements 118 a and diverter elements 112 acan be located at or near the proximal end of an implanted lead 104 ₁.These would be preferably located in an implanted lead proximal maleconnector shown as diverter element 112 a associated with one or moreimpeder elements 118 a. These impeder and diverter elements werepreviously described in FIGS. 2 through 11. There is an advantage tolocating the tuned diverter elements 112 outside of the AIMD housing.This can also be true for the impeder elements, but is not as critical.The reason has to do with the high volumes that certain AIMDs aremanufactured in. For example, there are over 650,000 cardiac pacemakersmanufactured every year. It would be highly impractical to have adifferent pacemaker design associated with every different type of lead.Hospitals, in general, inventory a great number of leads of varyinglengths. This is because they can be used in pediatric applications,children and full-size adult applications. Accordingly, implantedcardiac leads can vary anywhere from 20 cm to over 60 cm in length insome cases. The characteristic impedance of these implanted leads variesnot only with their length, but with their implant geometries andtrajectories through body tissues. In accordance with the tuned energybalance principles of the present invention, ideally, the reactivecomponent of the lead, which is usually an inductive reactance, would becanceled by the reactance of the impeder element 112 of the presentinvention (ref. FIGS. 89-93). As previously mentioned, since cardiacpacemakers are manufactured in high volume automated facilities, itsreally not practical to have a custom internal impeder element 112 foreach and every external implanted lead possibility. An additionalcomplication occurs because of the tendency to mix and match. That is,it is very common in medical practice to use one manufacturer'spacemaker with another manufacturer's leads. In other words, a St. Judepacemaker might be implanted with Medtronic leads. Installation of thecardiac pacemaker header block HB is a subsequent manufacturingoperation. Therefore, it would be relatively easy and inexpensive tocustom tailored diverter elements 112 a and impeder elements 118 a to aparticular lead. In a particularly preferred embodiment, the impeder 112and diverter 118 elements would be located at or near the proximal endof a lead 104 ₁ as shown. In this case, each lead could have its owncustom impeder element 112 b whose reactance would be tuned and equaland opposite to that of the characteristic reactance (impedance) of thelead. This would achieve optimal energy transfer through the wiring ofthe AIMD to the energy dissipating surface housing 161 in accordancewith the present invention. Referring once again to FIG. 115, one cansee that there is a novel switch 266 ₁ in accordance with the presentinvention that is shown inside the housing of the AIMD. This can, ofcourse, be mounted on an AIMD circuit board substrate or the like. Inthe MRI compatible mode, the switch grounds at least one of the leadwires that are routed to the proximal connector assembly PCA in order toprovide a ground connection to diverter element 112 a. This provides acircuit path from the one or more diverter elements located at or nearthe proximal end of lead 104 ₁ and/or in a proximal lead connector PCAsuch that high frequency energy picked up by the lead LW₁ can be routedthrough lead 118 to switch 266 ₁ and in turn to the energy dissipatingsurface 161. In this way, impeder element 112 a and diverter elements118 a work in accordance with the present invention.

FIG. 116 illustrates an alternative embodiment wherein the frequencydiverter element 112 is any of the previously shown or describeddiverter elements including short circuits, capacitors, L-C traps or lowpass filters and the like. The diverter element(s) would be permanentlyconnected to the AIMD housing 124 as shown. In a preferred embodiment,the diverter element 112 is frequency selective. A switch 266 ₁ shown inthe open position would be incorporated between the frequency diverterelement 112 and any non-linear AIMD electronics, such as those oncircuit board 137. In FIG. 116, the switch 266 ₁ is shown in the openposition which is also the MRI compatible mode for the AIMD. Duringnormal operation in the patient's normal daily environment, switch 266 ₁would normally be closed (not shown). This would connect the AIMDelectronics 137 to the implanted lead 104 ₁. In this normal operatingmode, the frequency selective diverter element 112 could not be a shortcircuit. In a preferred embodiment, it would be a frequency selectivereactance, such as a capacitor or an L-C trap filter. During this normalmode of operation, the AIMD would have EMI protection from diverterelement 112. In the MRI compatible mode, switch 266 ₁ is opened up asshown. In this case, the AIMD electronics would be completelydisconnected from the implanted lead 104 ₁. In general, this would workvery well for AIMDs that are not life-sustaining. For a spinal cordstimulator patient, urinary incontinence patient, or a cochlear implantpatient, it really is no more than a small inconvenience to have theAIMD inoperable during the relatively short time of an MRI scan.Immediately after the MRI scanning was complete, the device would bere-programmed or reset so that the switch 266 ₁ would be switched backto its closed position also known as the AIMD normal operating mode. Ifthe AIMD were a cardiac pacemaker and the patient were pacemakerdependent, then it would not be possible to disconnect the AIMDelectronics 137 from the implanted lead 104 ₁ (lack of pacing pulseswould be life threatening). In this case, alternative methods would berequired to protect the implanted lead from overheating and also protectAIMD electronics from EMI. In particular, its very important that theswitch 266 ₁ be incorporated between the frequency selective diverterelement (s) 112 and any non-linear circuit elements, such asover-voltage or transient protection diodes, Transorbs and the like.Non-linear circuit elements can act as detectors for pulse modulationthat could be coupled to implanted leads from the high powered MRI RF orgradient fields. As previously illustrated in FIG. 101 and described, itis important that RF rectification pulses not appear on the implantedlead where they could potentially capture (stimulate) or damage bodytissues at the electrode interface. In a preferred embodiment, thefrequency selective diverter element 112 could be a feedthroughcapacitor 114 (FIG. 39) which would be mounted directly on the hermeticterminal 128 of the AIMD. In this case, switch 266 ₁ would be disposedbetween the frequency diverter element 112 and all non-linear or otherAIMD internal electronic circuits. This would be a safety measure tomake sure that no non-linear circuit elements would be connected. Thisis very important so that there would be no chance for MRI RF orgradient rectification or tissue stimulation ratification to occurduring high intensity MRI scans.

Referring once again to FIG. 116, there are a number of very importantadvantages of this switched safety protection circuit for an AIMD systemduring exposure to high power electromagnetic fields such as thoseproduced by MRI scanners. The AIMD system includes an active implantablemedical device, such as a cardiac pacemaker, which generally includesinternal electronic circuits. The system also includes at least oneimplanted lead or leadwire associated with the AIMD electronics. Part ofthe system is an energy dissipating surface 161, which in a preferredembodiment, is the conductive AIMD housing 124. As illustrated, one ormore diversion circuits 112 are associated with the energy dissipatingsurface 161. In general, the diversion circuit is coupled between theimplanted lead or leadwire and the AIMD housing 124 which acts as anenergy dissipating surface 161. In a particularly preferred embodiment,the diversion circuit is a frequency selective diversion circuit, whichincludes reactances. As can be seen, at least one switch is disposedbetween the diversion circuit and the AIMD electronics. The purpose ofthe switch is to open up or disconnect the AIMD electronics from theimplanted lead while at the same time diverting energy in the implantedlead or leadwire to the energy dissipating surface 161. It will beobvious to those skilled in the art that switch 266 can be a singlepole-single throw or a multi-pole single throw. It is common for AIMDsto have not just one (unipolar), but a number of implanted leads. Inaccordance with the present invention, switch 266 ₁ as described in FIG.116 could be disposed in one, several or even all of the implanted leador leadwire circuits (multi-pole). The switch circuit of FIG. 116accomplishes three very important objectives: (1) it disconnects AIMDelectronics therefore making them highly immune to EMI that may becoupled onto leads in the presence of high power electromagnetic fields.High powered electromagnetic fields include the RF pulsed and gradientfields produced by an MRI scanner; (2) the switch 266 ₁ also disconnectsany non-linear AIMD circuit elements. This prevents a condition known asRF or gradient field rectification. As previously described, RF orgradient rectification can cause low frequency pulses to appear onimplanted leads which could improperly stimulate or damage body tissues.For example, low frequency pulses at 300 Hz may cause the heart to beattoo fast and lead to a very dangerous condition known as ventricularfibrillation; (3) the permanently wired diverter 112 of FIG. 116 acts topull unwanted RF energy from the implanted lead and divert it to theAIMD housing 124 which acts as an energy dissipating surface 161. Thisis very important to prevent overheating of the implanted lead and/orits distal electrode-to-tissue interface.

FIG. 117 is very similar to FIG. 116 except that the switch 266 ₂ hasbeen reversed and is now a single pole double throw switch. It operatesvery much as described in FIG. 116, but it also offers anotheradvantage. When the switch is switched open into the MRI compatibleposition as shown, the wiper is also grounded to the housing 124 of theAIMD. This is an additional safety feature. By grounding the AIMDelectronics 137, one provides additional protection against the chancethat EMI could cross couple over to sensitive AIMD circuitry. Aspreviously mentioned, switch 266 ₂ is shown in the MRI compatibilityposition. As previously described in FIG. 116, the switch would beswitched into the closed position during the normal AIMD operating mode.In other words, during daily life, the patient's AIMD electronics 137would be directly connected through switch 266 ₂ to the implanted lead104 ₁.

In summary, compatibility of probes, catheters, cardiac pacemakers,implantable defibrillators and other types of active implantable medicaldevices with magnetic resonance imaging and other types of hospitaldiagnostic equipment has become a major issue. The present inventionaddresses this by providing an overall energy management system which iscapable of controlling the energy induced in implanted leads from RFpulsed fields, such as those generated by MRI scanners. Moreparticularly, a tuned energy balanced system including a switched safetyprotection circuit minimizes heating of an implanted lead in a highpower electromagnetic field environment. The tuned energy balancedsystem comprises an implanted lead having impedance characteristics at aselected RF frequency or frequency band, and an energy dissipatingsurface associated with the implanted lead. A diversion circuitconductively couples the implanted lead to the energy dissipatingsurface. The diversion circuit comprises one or more passive electronicnetwork components whose impedance characteristics are at leastpartially tuned to the implanted lead's impedance characteristics, tofacilitate transfer to the energy dissipating surface of high frequencyenergy induced on the implanted lead at the selected RF frequency orfrequency band.

The switched safety protection circuit includes at least one switchdisposed between the diversion circuit and the AIMD electronics fordiverting energy in the implanted lead or the leadwire through thediversion circuit to the energy dissipating surface. The switch isdisposed so as to electrically open the implanted lead or the leadwirewhen diverting energy in the implanted lead or the leadwire through thediversion circuit to the energy dissipating surface. Alternatively or inaddition, a switch may be disposed between the implanted lead or theleadwire and the diversion circuit. In preferred embodiments, theswitches may comprise a single or multi-pole double throw switch, or asingle or multi-pole single throw switch.

The tuned energy balanced system utilizing the switched safetyprotection circuit of the present invention harmlessly shunts RF energyinduced on an implanted lead or leadwire into either an EDS surface, theAIMD housing, a bulk thermal mass, or the handle of a probe or catheter.

Impedance circuits may be combined with the diversion or divertercircuits to raise and further control the overall impedance of thesystem to achieve maximal energy transfer and minimum thermal rise inthe implanted lead system.

Although several embodiments have been described in detail for purposesof illustration, various modifications may be made without departingfrom the scope and spirit of the invention. Accordingly, the inventionis not to be limited, except as by the appended claims.

What is claimed is:
 1. An implantable medical device, comprising: a) athermally or electrically conductive housing for the implantable medicaldevice containing tissue-stimulating or biological-sensing circuits; b)a feedthrough terminal disposed in the conductive housing; c) a firstconductive path electrically coupled between the feedthrough terminaland the tissue-stimulating or biological-sensing circuits; d) a switchelectrically coupled in series along the first conductive path, theswitch comprising a first throw end, a second throw end and a pole end,wherein the pole end is permanently electrically coupled to thetissue-stimulating or biological-sensing circuit and the first throw endis permanently connected to the feedthrough terminal; and e) a secondconductive path permanently electrically coupled between the secondthrow end of the switch and the conductive housing; f) wherein theswitch is configured to be selectively actuatable in response to aprogrammable telemetry signal to either connect the pole to the firstthrow end or to connect the pole to the second throw end.
 2. Theimplantable medical device of claim 1, wherein in a normal operatingmode the pole is connected to the first throw end.
 3. The implantablemedical device of claim 2, wherein in an MRI mode the pole is connectedto the second throw end.
 4. The implantable medical device of claim 1,including a third conductive path coupled at a first end anywhere alongthe first conductive path between the feedthrough terminal and the firstthrow end of the switch, and coupled at a second end to the conductivehousing, wherein a frequency variable diverter element is coupled inseries along the third conductive path.
 5. The implantable medicaldevice of claim 4, wherein the frequency variable diverter elementcomprises a low pass filter, a capacitor, a feedthrough capacitor, an LCtrap filter, a Pi filter, a T filter, an LL filter or an “n” elementfilter.
 6. The implantable medical device of claim 1, wherein the switchcomprises a multi-pole double throw switch.
 7. The implantable medicaldevice of claim 1, wherein the switch comprises a MEMS switch, amechanical switch, a reed switch, an electronic switch, a programmableswitch, an automatically actuated switch, a Hall Effect switch, a fieldeffect transistor (FET) switch or a PIN diode.
 8. The implantablemedical device of claim 1, wherein the tissue-stimulating orbiological-sensing circuits comprise non-linear circuit elements,over-voltage circuit elements, transient voltage suppression diodes,Transorbs, AIMD sensing circuits or active implantable medical devicetherapy delivery circuits.
 9. The implantable medical device of claim 4,wherein the frequency variable diverter element comprises a high passfilter.
 10. The implantable medical device of claim 9, wherein the highpass filter comprises a capacitor or a resistor in series with acapacitor.
 11. The implantable medical device of claim 4, wherein thefrequency variable diverter element comprises a plurality of LC trapfilters, each LC trap filter being resonant at a different MRI RF pulsedfrequency.
 12. The implantable medical device of claim 1, wherein theimplantable medical device comprises an implantable hearing device, aneurostimulator, a brain stimulator, a cardiac pacemaker, a leftventricular assist device, an artificial heart, a drug pump, a bonegrowth stimulator, a urinary incontinence device, a spinal, cordstimulator, an anti-tremor stimulator, an implantable cardioverterdefibrillator, a congestive heart failure device, a probe, a catheter ora cardio resynchronization therapy device.
 13. The implantable medicaldevice of claim 4, wherein the third conductive path is disposed withinthe conductive housing.
 14. The implantable medical device of claim 4,wherein third conductive path is disposed within a header block which isattached to the conductive housing.
 15. The implantable medical deviceof claim 14, including an EMI shield conductively coupled to theconductive housing and coaxially extending about the third conductivepath in non-conductive relation thereto.
 16. An implantable medicaldevice, comprising: a) a thermally or electrically conductive housingfor the implantable medical device containing tissue-stimulating orbiological-sensing circuits; b) a feedthrough terminal disposed in theconductive housing; a first conductive path electrically coupled betweenthe feedthrough terminal and the tissue-stimulating orbiological-sensing circuits; d) a switch electrically coupled in seriesalong the first conductive path, the switch comprising a first throwend, a second throw end and a pole end, wherein the pole end ispermanently electrically coupled to the tissue-stimulating orbiological-sensing circuit and the first throw end is permanentlyconnected to the feedthrough terminal; and e) a second conductive pathpermanently electrically coupled between the second throw end of theswitch and the conductive housing; f) wherein the switch is configuredto be selectively actuatable in response to a programmable telemetrysignal to either connect the pole to the first throw end in a normaloperating mode or to connect the pole to the second throw end in an MRImode; and g) a third conductive path coupled at a first end anywherealong the first conductive path between the feedthrough terminal and thefirst throw end of the switch, and coupled at a second end to theconductive housing, wherein a frequency variable diverter element iscoupled in series along the third conductive path.
 17. An implantablemedical device, comprising: a) a thermally or electrically conductivehousing for the implantable medical device containing tissue stimulatingor biological-sensing circuits; b) a feedthrough terminal disposed inthe conductive housing; c) a first conductive path electrically coupledbetween the feedthrough terminal and the tissue-stimulating orbiological-sensing circuits; d) a switch electrically coupled in seriesalong the first conductive path, the switch comprising a first throwend, a second throw end and a pole end, wherein the pole end ispermanently electrically coupled to the tissue-stimulating orbiological-sensing circuit and the first throw end is permanentlyconnected to the feedthrough terminal; and e) a second conductive pathpermanently electrically coupled between the second throw end of theswitch and the conductive housing; f) wherein the switch is configuredto be actuatable in response to a magnetic resonance B₀ field sensorthat, in the absence of a static magnetic field, electrically connectsthe pole to the first throw end, and in the presence of a staticmagnetic field, electrically connects the pole to the second throw end.